Systems for Comprehensive Fourier Domain Optical Coherence Tomography (FDOCT) and Related Methods

ABSTRACT

Optical coherence tomography systems for imaging a whole eye are provided including a sample arm including focal optics that are configured to rapidly switch between at least two scanning modes in less than about 1.0 second.

CLAIM OF PRIORITY

The present application is a continuation of U.S. application Ser. No.12/910,184, filed Oct. 22, 2010, which claims priority from U.S.Provisional Application No. 61/254,465, filed Oct. 23, 2009 and to U.S.patent application Ser. No. 12/887,891, filed on Sep. 22, 2010, now U.S.Pat. No. 8,348,427, the disclosures of which are hereby incorporatedherein by reference as if set forth in their entirety.

STATEMENT OF GOVERNMENT SUPPORT

This invention was made with government support under grant numbers2R44EY015585 and 2R43EY018021 awarded by National Institutes of Health,National Eye Institute. The United States Government has certain rightsin this invention.

FIELD

The present inventive concept generally relates to imaging and, moreparticularly, to frequency domain optical coherence tomography (FDOCT)and related systems and methods.

BACKGROUND

Optical coherence tomography (OCT) is a noninvasive imaging techniquethat provides microscopic tomographic sectioning of biological samples.By measuring singly backscattered light as a function of depth, OCTfills a valuable niche in imaging of tissue ultrastructure, providingsubsurface imaging with high spatial resolution (˜2.0-10.0 μm) in threedimensions and high sensitivity (>110 dB) in vivo with no contact neededbetween the probe and the tissue.

In biological and biomedical imaging applications, OCT allows formicrometer-scale imaging non-invasively in transparent, translucent,and/or highly-scattering biological tissues. The longitudinal rangingcapability of OCT is generally based on low-coherence interferometry, inwhich light from a broadband source is split between illuminating thesample of interest and a reference path. The interference pattern oflight reflected or backscattered from the sample and light from thereference delay contains information about the location and scatteringamplitude of the scatterers in the sample. In time-domain OCT (TDOCT),this information is typically extracted by scanning the reference pathdelay and detecting the resulting interferogram pattern as a function ofthat delay. The envelope of the interferogram pattern thus detectedrepresents a map of the reflectivity of the sample versus depth,generally called an A-scan, with depth resolution given by the coherencelength of the source. In OCT systems, multiple A-scans are typicallyacquired while the sample beam is scanned laterally across the tissuesurface, building up a two-dimensional map of reflectivity versus depthand lateral extent typically called a B-scan. The lateral resolution ofthe B-scan is approximated by the confocal resolving power of the samplearm optical system, which is usually given by the size of the focusedoptical spot in the tissue.

The time-domain approach used in conventional OCT, including commercialinstruments, such as Carl Zeiss Meditec's Stratus® and Visante®products, has been successful in supporting biological and medicalapplications, and numerous in vivo human clinical trials of OCT reportedto date have utilized this approach.

An alternate approach to data collection in OCT has been shown to havesignificant advantages in increased signal-to-noise ratio (SNR). Thisapproach involves acquiring the interferometric signal generated bymixing sample light with reference light at a fixed group delay as afunction of optical wavenumber. Two distinct methods have been developedwhich use this Fourier domain OCT (FD-OCT) approach. The first,generally termed Spectral-domain or spectrometer-based OCT (SDOCT), usesa broadband light source and achieves spectral discrimination with adispersive spectrometer in the detector arm. The second, generallytermed swept-source OCT (SSOCT) or optical frequency-domain imaging(OFDI), time-encodes wavenumber by rapidly tuning a narrowband sourcethrough a broad optical bandwidth. Both of these techniques may allowfor a dramatic improvement in SNR of up to 15.0-20.0 dB over time-domainOCT, because they typically capture the A-scan data in parallel. This isin contrast to previous-generation time-domain OCT, where destructiveinterference is typically used to isolate the interferometric signalfrom only one depth at a time as the reference delay is scanned.

FDOCT systems are discussed below with respect to FIGS. 1 through 3.Referring first to FIG. 1, a block diagram illustrating a Fourier domainOCT system in accordance with some embodiments of the present inventiveconcept will be discussed. As illustrated in FIG. 1, the system includesa broadband source 100, a reference arm 110 and a sample arm 140 coupledto each other by a beamsplitter 120. The beamsplitter 120 may be, forexample, a fiber optic coupler or a bulk or micro-optic coupler withoutdeparting from the scope of the present inventive concept. Thebeamsplitter 120 may provide from about a 50/50 to about a 90/10 splitratio. As further illustrated in FIG. 1, the beamsplitter 120 is alsocoupled to a wavelength or frequency sampled detection module 130 over adetection path 106 that may be provided by an optical fiber.

As further illustrated in FIG. 1, the source 100 is coupled to thebeamsplitter 120 by a source path 105. The source 100 may be, forexample, a SLED or tunable source. The reference arm 110 is coupled tothe beamsplitter over a reference arm path 107. Similarly, the samplearm 140 is coupled to the beamsplitter 120 over the sample arm path 108.The source path 105, the reference arm path 107 and the sample arm path108 may all be provided by optical fiber.

As further illustrated in FIG. 1, the sample arm 140 may includescanning delivery optics and focal optics 160. Also illustrated in FIG.1 is the reference plane 150 and a representation of an OCT imagingwindow 170.

Referring now to FIG. 2, a block diagram of an FDOCT retinal imagingsystem will be discussed. As illustrated in FIG. 2, in an FDOCT retinalimaging system, the reference arm 110 may further include a collimatorassembly 280, a variable attenuator 281 that can be neutral density orvariable aperture, a mirror assembly 282, a reference arm variable pathlength adjustment 283 and a path length matching position 250, i.e.optical path length reference to sample. As further illustrated, thesample arm 240 may include a dual-axis scanner assembly 290 and avariable focus objective lens 291.

The sample in FIG. 2 is an eye including a cornea 295, iris/pupil 294,ocular lens 293 and retina 296. A representation of an OCT imagingwindow 270 is illustrated near the retina 296. The retinal imagingsystem relies in the optics of the subject eye, notably cornea 295 andocular lens 293, to image the posterior structures of the eye.

Referring now to FIG. 3A, a block diagram illustrating a FDOCT corneaimaging system will be discussed. As illustrated therein, the system ofFIG. 3A is very similar to the system of FIG. 2. However, the objectivelens variable focus need not be included, and is not included in FIG.3A. The anterior imaging system of FIG. 3A images the anteriorstructures directly, without reliance on the optics of the subject tofocus on the anterior structures.

As illustrated by FIGS. 3A through 3C, the OCT imaging window 370 can bemoved to image various portions of the sample.

In both spectrometer-based and swept-source implementations of FDOCT,light returning from all depths is generally collected simultaneously,and is manifested as modulations in the detected spectrum.Transformation of the detected spectrum from wavelength to wavenumber(or frequency), correction for dispersion mismatches between the sampleand reference arms, and Fast Fourier transformation typically providesthe spatial domain signal or “A-scan” representing depth-resolvedreflectivity of the sample. The uncorrected A-scan may also include astrong DC component at zero pathlength offset, so-called“autocorrelation” artifacts resulting from mutual interference betweeninternal sample reflections, as well as both positive and negativefrequency components of the depth-dependent cosine frequencyinterference terms. Because of this, FDOCT systems typically exhibit a“complex conjugate artifact” due to the fact that the Fourier transformof a real signal, the detected spectral interferogram, is typicallyHermitian symmetric, i.e., positive and negative spatial frequencies arenot independent. As a consequence, sample reflections at a positivedisplacement, relative to the reference delay, typically cannot bedistinguished from reflections at the same negative displacement, andappear as upside-down, overlapping images on top of genuine samplestructure, which generally cannot be removed by image processing.

The maximum single-sided imaging depth available in SDOCT is governed bythe spectral sampling interval. The maximum single-sided imaging depthis inversely proportional to the spectral sampling interval. With afixed number of sampled spectral elements, there is an inverserelationship between the maximum imaging depth and the minimum axialresolution of the imaging system. In commercial FDOCT systems at 830 nmand 1300 nm reported to date, the single-sided imaging depth has beenlimited to approximately 4 mm. Time domain imaging has been used forgreater imaging depths.

The finite spectral resolution of any real FDOCT system, whethergoverned by the linewidth of a swept laser source in SSOCT, or thegeometric optical performance of the spectrometer convolved with thefinite pixel size of the detector array in SDOCT, gives rise to asensitivity “falloff” with imaging depth into the sample. It is commonto have greater than 6 dB degradation in signal-to-noise from theposition of zero reference delay to the position of maximum single-sideddepth. This sensitivity “falloff” limits the portion of the single-sideddepth useful for imaging.

To reduce the impact of these limitations in FDOCT imaging, imaging iscommonly performed with the entire sample either above or below thereference position, limiting the available imaging depth to between 2 mmand 4 mm, and placing the sample region of interest close to the zeroreference delay position.

Each of these constraints poses limitations on the application of FDOCTto clinical ophthalmology. Imaging systems have generally been dedicatedto imaging of specific anatomy, such as retina or cornea, where themirror image artifacts do not fold over onto images of the region ofinterest. Utility to image deeper anatomic structures, such as thechoroid, has been limited by sensitivity “falloff”.

Addressing these limitations opens significant new application areas forFDOCT, particularly in ophthalmology. Full range volumetric anteriorsegment imaging (cornea to lens) for improved diagnosis of narrow angleglaucoma is enabled at speeds 20 times greater and resolutions fourtimes finer than time domain implementations. Real-time image guidedsurgery, for anterior chamber, cataract, or retina, is enabled byallowing placement of a deep imaging window at any position within thesample, without concern for confounding mirror image artifacts or signal“falloff,” Images of the entire eye may be acquired, enabling for thefirst time modeling in three dimensions the entire optical structure ofthe eye and enabling whole-eye biometry.

SUMMARY

Some embodiments discussed herein provide an optical coherencetomography system for imaging a whole eye, the system includes a samplearm including focal optics that are configured to rapidly switch betweenat least two scanning modes in less than about 1.0 second.

In further embodiments, the focal optics may be configured to beswitched between the at least two modes without use of an externaladapter.

In still further embodiments, the at least two modes may include ananterior segment scanning mode and a retinal scanning mode. The systemmay further include a mechanical means configured to rapidly insert atleast one additional lens into and/or remove the at least one additionallens from an optical path of the sample arm to switch the system betweenthe anterior segment scanning mode and the retinal scanning mode.

In some embodiments, the sample arm of the system in retinal scanningmode may include a collimator, a two-dimensional galvanometer scanner,and a single scan lens in a telecentric configuration. The mechanicalmeans may be configured to rapidly insert a single additional lens intothe optical path immediately proximal or immediately distal to thecollimating lens to change the system from the retinal scanning mode toanterior segment scanning mode.

In further embodiments, the additional lens in the optical path may beconfigured to change a sample arm beam from collimated to focusing onthe two-dimensional galvanometer scanner.

In still further embodiments, the mechanical means includes at least onelens mounted to a mechanical plate that is configured to be rotated intoand out of the optical path. The system may further include a controllerconfigured to cause the mechanical means to rapidly rotate the plate.

In some embodiments, the mechanical means may include a rotary solenoidattached to an arm including the additional lens, the rotary solenoidmay be configured to rapidly rotate the additional lens into and out ofthe optical path.

In further embodiments, the at least two modes may include an anteriorsegment scanning mode and a retinal scanning mode. The sample arm of thesystem in the retinal scanning mode may include a collimating lens, atwo-dimensional galvanometer scanner pair, a scan lens and an objectivelens. The sample arm of the system in the anterior segment scanning modemay include a collimating lens, two two-dimensional galvanometer scannerpairs, a scan lens, an objective lens and a curved mirror placed a focallength f away from a first of the two two-dimensional galvanometerscanner pairs, wherein the a second of the two two-dimensionalgalvanometer scanner pairs directs re-directed collimated light in atriangular pattern towards the curved mirror causing an optical pathlength of the system to be longer in the anterior segment scanning modeas compared to the retinal scanning mode.

In still further embodiments, the at least two modes may include ananterior segment scanning mode and a retinal scanning mode. The samplearm of the system in the retinal scanning mode may include a collimatinglens, a two-dimensional galvanometer scanner pair, a scan lens and anobjective lens. The sample arm of the system in the anterior segmentscanning mode may include a collimating lens, a two-dimensionalgalvanometer scanner pair, a scan lens, an objective lens, a flat mirrorand a concave mirror placed a focal length f away from thetwo-dimensional galvanometer scanner pairs, wherein light incident onthe two-dimensional scanner pair is deviated such that the an incidentcollimated beam is directed into a separate path consisting of the flatmirror and the concave mirror.

In some embodiments, the system may further include a reference armincluding a means for rapidly switching a reference delay when the focaloptics of the sample arm are switched between the at least two scanningmodes. The reference arm may further include a coupler configured tosplit light from the reference arm of the interferometer into at leasttwo separate paths. The at least two separate paths may be preset to anoptical delay each corresponding to one of the at least two scanningmodes.

In further embodiments, the means for rapidly switching a referencedelay may include a rapid mechanical switch configured to block all buta desired reference delay associated with a corresponding one of the atleast two scanning modes.

In still further embodiments, the at least two scanning modes mayinclude and iris pivot scanning mode and a telecentric scanning mode.The sample arm may include a telecentric scanning lens and first andsecond objective lenses, wherein the first and second objective lensesare a first distance apart in a first position when the system isoperating in telecentric scanning mode and a second distance apart in asecond position when the system is operating in the iris pivot scanningmode.

In some embodiments, the first and second objective lenses may beconfigured to slide between the first and second positions to switchbetween scanning modes.

In further embodiments, the at least two scanning modes may include atelecentric scanning mode and a collimated scanning mode. The sample armmay include a fiber input, a collimating lens, a scanning mirror, atelecentric scanning lens, and a telecentric scanning beam. Thecollimating lens may be in a first position in telecentric scanning modeand a second position in collimated scanning mode. The collimating lensmay be translated by a distance equal to a focal length of thecollimating lens in the second position to provide the collimatedscanning mode.

In still further embodiments, the at least two scanning modes mayinclude a telecentric scanning mode and a collimated scanning mode. Thesample arm may include a fiber input, a collimating lens, a scanningmirror, a telecentric scanning lens, and a telecentric scanning beam inthe telecentric scanning mode. The sample arm may further include asecondary lens behind the collimating lens in collimated scanning mode.

Some embodiments of the present invention provide optical coherencetomography systems for imaging a whole eye, the system comprising areference arm configured to adapt to focal optics of at least twoscanning modes of the system.

In further embodiments, the reference arm may include a mechanical meansconfigured to discretely switch reference arms such that the referencearm is matched a corresponding one of the at least two scanning modes.

In still further embodiments, the reference arm may be configured torapidly switch between reference delays, each of the reference delayscorresponding to one of the at least two scanning modes.

Some embodiments provide methods for imaging a whole eye in an opticalcoherence tomography system, the methods including rapidly switchingfocal optics of a sample arm between at least two scanning modes in lessthan about 1.0 second.

In further embodiments, switching the focal optics of the sample arm mayinclude switching the focal optics of the sample arm between the atleast two scanning modes without use of an external adapter.

In still further embodiments, the at least two modes may include ananterior segment scanning mode and a retinal scanning mode.

In some embodiments, the method may further include rapidly switching areference delay when the focal optics of the sample arm are switchedbetween the at least two scanning modes.

BRIEF DESCRIPTION OF THE FIGURES

FIG. 1 is a block diagram illustrating a Fourier domain opticalcoherence tomography (OCT) imaging system.

FIG. 2 is a block diagram illustrating a Fourier domain retinal opticalcoherence tomography system in accordance with some embodiments of theinventive concept.

FIGS. 3A through 3C are block diagrams illustrating a Fourier domaincorneal optical coherence tomography system in accordance with someembodiments.

FIGS. 4A and 4B are graphs illustrating the spatial and spectral domain,respectively, that illustrate conceptually the effects that variousparameters of the light source and spectral detection system inaccordance with some embodiments.

FIGS. 5A through 5E OCT systems and images acquired using these OCTophthalmic imaging systems in accordance with some embodiments.

FIG. 6 is a block diagram illustrating an FDOCT interferometer having avariable round trip phase delay in the reference arm in accordance withsome embodiments.

FIG. 7 is a block diagram illustrating a GRISM in accordance with someembodiments of the present inventive concept.

FIG. 8 is a graph illustrating grating and prism dispersion inaccordance with some embodiments of the present inventive concept.

FIG. 9 is a graph illustrating cumulative wavenumber sampling shift as acumulative error in channel position from pixel position, as referencedto the center pixel, due to dispersion in accordance with someembodiments of the inventive concept.

FIG. 10 is a graph illustrating GRISM angle dispersion in accordancewith some embodiments of the present inventive concept.

FIG. 11 is a schematic block diagram illustrating an optical coherencetomography (OCT) system including a piezoelectric transducer (PZT)element in accordance with some embodiments of the present inventiveconcept.

FIGS. 12 and 13A through 13E are diagrams illustrating a series ofimaging windows that may be applied for a select variety of imagingcircumstances in accordance with some embodiments of the inventiveconcept.

FIG. 14 is a block diagram of an extended depth fourier domain OCTimaging system in accordance with some embodiments of the inventiveconcept.

FIG. 15 is a graph illustrating depth and resolution vs. spectrometerbandwidth and samples for an extended depth FDOCT system in accordancewith some embodiments discussed herein.

FIG. 16 is a graph illustrating image depth and sampling free spectralrange vs. spectrometer bandwidth for an extended depth FDOCT system inaccordance with some embodiments of the present inventive concept.

FIG. 17 is a table including various details with respect to SDOCTsystems in accordance with various embodiments of the present inventiveconcept.

FIGS. 18A through 18C are block diagrams illustrating variousembodiments of extended-depth FDOCT imaging systems in accordance withsome embodiments of the inventive concept.

FIG. 19 is a block diagram of an FDOCT system including a swept sourceand optical filter in accordance with some embodiments.

FIG. 20 is a block diagram illustrating an optical filter configurationin accordance with some embodiments.

FIGS. 21 through 24 are graphs illustrating various aspects of output ofthe optical filter of FIG. 20 in accordance with some embodiments.

FIG. 25 is a block diagram illustrating data flow of an SDOCT system inaccordance with some embodiments.

FIG. 26 is a complex conjugate removal (CCR) control timing diagram inaccordance with some embodiments.

FIGS. 27A through 27C illustrate diagrams in accordance with conventionanterior segment and retinal OCT sample arm scanning.

FIGS. 28A and 28B illustrate switchable anterior-retinal scanningsystems in accordance with some embodiments.

FIGS. 29A and 29B illustrate switchable anterior-retinal scanningsystems in accordance with some embodiments.

FIGS. 30A and 30B illustrate switchable anterior-retinal scanningsystems in accordance with some embodiments.

FIG. 31 is a diagram illustrating a rapidly switching reference delay inaccordance with embodiments illustrated in FIGS. 28A through 30B.

FIG. 32 is a block diagram illustrating a comprehensive Ocular SpectralDomain OCT imaging system in accordance with some embodiments.

FIG. 33 is a block diagram illustrating a comprehensive ocular sweptsource OCT imaging system in accordance with some embodiments.

FIG. 34 is a diagram illustrating an optical layout for telecentricscanning mode in accordance with some embodiments.

FIG. 35 is a diagram illustrating an optical system in concentric orretinal scanning mode in accordance with some embodiments.

FIG. 36 is a diagram illustrating an optical layout for telecentricscanning mode in accordance with some embodiments.

FIG. 37 is a diagram illustrating an optical layout for collimated modein accordance with some embodiments.

FIG. 38 is a diagram illustrating an optical layout for collimated modein accordance with some embodiments.

FIGS. 39 through 41 are flowcharts illustrating scanning methods inaccordance with various embodiments discussed herein.

FIG. 42 is a block diagram illustrating a mechanical means for rapidlyswitching scanning modes in accordance with some embodiments of theinventive concept.

DETAILED DESCRIPTION OF EMBODIMENTS

Specific exemplary embodiments of the inventive concept now will bedescribed with reference to the accompanying drawings. This inventiveconcept may, however, be embodied in many different forms and should notbe construed as limited to the embodiments set forth herein; rather,these embodiments are provided so that this disclosure will be thoroughand complete, and will fully convey the scope of the inventive conceptto those skilled in the art. The terminology used in the detaileddescription of the particular exemplary embodiments illustrated in theaccompanying drawings is not intended to be limiting of the inventiveconcept. In the drawings, like numbers refer to like elements.

As used herein, the singular forms “a”, “an” and “the” are intended toinclude the plural forms as well, unless expressly stated otherwise. Itwill be further understood that the terms “includes,” “comprises,”“including” and/or “comprising,” when used in this specification,specify the presence of stated features, integers, steps, operations,elements, and/or components, but do not preclude the presence oraddition of one or more other features, integers, steps, operations,elements, components, and/or groups thereof. It will be understood thatwhen an element is referred to as being “connected” or “coupled” toanother element, it can be directly connected or coupled to the otherelement or intervening elements may be present. Furthermore, “connected”or “coupled” as used herein may include wirelessly connected or coupled.As used herein, the term “and/or” includes any and all combinations ofone or more of the associated listed items.

Unless otherwise defined, all terms (including technical and scientificterms) used herein have the same meaning as commonly understood by oneof ordinary skill in the art to which this inventive concept belongs. Itwill be further understood that terms, such as those defined in commonlyused dictionaries, should be interpreted as having a meaning that isconsistent with their meaning in the context of the relevant art andthis specification and will not be interpreted in an idealized or overlyformal sense unless expressly so defined herein.

Embodiments discussed herein with respect to FIGS. 1 through 26 may beused in combination with many imaging systems for many applications andis not limited to the specific systems or applications discussed herein.For example, imaging systems discussed in commonly assigned U.S. Pat.Nos. 7,602,500; 7,742,174; and 7,719,692 and U.S. Patent ApplicationPublications Nos. 2008/01606696 and 2008/0181477 may be used incombination with embodiments discussed herein. The disclosures of thesepatents and publications are hereby incorporated herein by reference asif set forth in its entirety.

As used herein, the term “spectral element” refers to the individuallyresolved samples of the interferometric spectrum as they are detected inan FDOCT system, the detected set of which forms the input to theFourier transform operation; a spectral element is characterized by afinite wavelength range that is generally a continuous small fraction ofthe total bandwidth, an optical power, and a power spectral density(lineshape).

The first successful clinical application of OCT was for high-resolutionimaging of ocular structure. OCT is well suited to ophthalmology becauseit is non-contact, easily adaptable to existing ophthalmicinstrumentation, and most importantly, the axial imaging resolution isindependent of the working distance. In the anterior eye, themicron-scale resolution of OCT imaging permits accurate biometry oflarge scale ocular structures and the evaluation of morphologicalchanges associated with pathologies of the cornea, iris, and lens. Inthe retina, OCT is the only technique capable of resolving retinalsubstructure in cross section in the living eye. Imaging of retinalsubstructure is clinically relevant to the diagnosis and management ofmany ocular diseases. In many clinical trials of OCT, striking imageshave been obtained of a variety of retinal abnormalities, includingmacular defects and retinal nerve fiber atrophy. Retinal OCT has becomewell accepted as a clinical adjunct to conventional macular photography,as well as a very popular research tool.

As is well known in the art, Fourier domain optical coherence tomography(FDOCT) has become the standard of care in clinical ophthalmology forimaging of the retina. The theory and practice of FDOCT is well knownand documented.

Several academic research groups have published detailed treatments ofthe dramatic 20-30 dB signal-to-noise predicted and actual performanceimprovement in FDOCT as compared to its time-domain counterparts.Despite this tremendous improvement, there are performance limitationsin FDOCT which do not have analogs in previous time-domain systems. Inboth spectrometer-based (SDOCT) and swept-source (SSOCT) implementationsof FDOCT, the wavenumber (k) resolved receiver current can berepresented by equation (1) set out below:

i(k)=½ρS(k)[R _(R) +R _(S)+2√{square root over (R _(R) R _(S))} cos(2kz₀+φ)]  (1)

where ρ is the receiver element responsivity, S(k) is the light sourcespectrum, R_(R) and R_(S) are the received power from the reference andsample arms, respectively, z₀ is the pathlength difference between thereference delay and a target reflection in the sample (2z_(o) is theround-trip pathlength difference), and φ is the phase offset of theinterferometer at zero pathlength optical delay. As in time domain OCT,the axial imaging resolution Δz is defined by the source centerwavelength λ₀ and FWHM bandwidth Δλ as shown in equation (2) set outbelow:

$\begin{matrix}{{\Delta \; z} = {\frac{2{\ln (2)}}{\pi}\frac{\lambda_{0}^{2}}{\Delta \; \lambda}}} & (2)\end{matrix}$

In FDOCT, light returning from all depths is collected simultaneously,and is manifested as modulations in the detected spectrum.Transformation of the detected spectrum from wavelength to wavenumber,correction for dispersion mismatches between the sample and referencearms, and Fast Fourier transformation provides the spatial domain signalor “A-scan” representing depth-resolved reflectivity of the sample.

FIGS. 4A and 4B illustrate conceptually the effects that variousparameters of the light source and spectral detection system have on thetransformed OCT signal. In both SDOCT and SSOCT systems, the detectedspectral signal is composed of a DC term and cosinusoidal terms withdepth-dependent frequency. This signal is enveloped by the sourcespectrum and convolved with the spectral resolution (δ_(r)k) of theFDOCT system. In SSOCT, the spectral resolution δ_(r)k is limited by theinstantaneous lineshape of the swept laser source, while in SDOCT,δ_(r)k is the spectral resolution of the spectrometer (which in turn maybe described as a convolution of the geometric optic spectrometerresolution with the detector array pixel dimensions). The detectedspectrum is sampled with spectral sampling interval δ_(r)k into Nspectral channels, each of these channels comprising a unique spectralelement. The Fourier transform of the detected spectral signal, i.e. theset of spectral elements, includes a strong DC component at zeropathlength offset, as well as both positive and negative frequencycomponents of the depth-dependent cosine frequency terms located atpositions ±2x₀. The shape of each peak is defined by the coherencefunction of the source, given by the inverse Fourier transform of itstotal detected power spectral density; as in time-domain OCT, the axialimaging resolution is inversely proportional to the total detectedspectral bandwidth of the light source.

There are at least three important limitations which are novel in FDOCT:complex conjugate artifact, maximum imaging depth, and sensitivityfalloff. The presence in the spatial-domain A-scan data of both positiveand negative frequency components of the spectral interferometric signalgives rise to the so-called “complex conjugate artifact” which typicallyrequires careful sample positioning to assure that overlapping negativefrequency components do not interfere with the principalpositive-frequency image as illustrated in FIG. 5C. Methods to reducethis “complex conjugate artifact” are now known in the art. Inparticular, FIG. 5A illustrates a 3×3 SDOCT system utilizing twospectrometers; FIG. 5B illustrates a 3×3 SSOCT system implementing DCsignal subtraction with balanced photodiode detectors D1 and D2; FIGS.5C and 5D illustrate unprocessed and complex conjugate resolved imagesof human anterior segment acquired in vivo with 3×3 SSOCT, using a firstimage processing processing method; and FIG. 5E illustrates improvedcomplex conjugate artifact removal (better than 30 dB) obtained using aquadrature projection algorithm.

The complex conjugate artifact in FDOCT may also be removed by utilizingprinciples and techniques related to phase-shift interferometry. If theinterferometer is modified to provide for the introduction of a variablesingle-pass phase delay φ (round-trip phase delay 2φ) between thereference and sample arms, then a set of spectral interferograms may beacquired with different phase delays which can be combined in signalprocessing to eliminate the undesired artifacts. For example, FIG. 6illustrates an FDOCT interferometer with a variable phase modulatorplaced in the reference arm, such that the reference field returningfrom the reference arm is modified.

Various technical solutions to this complex conjugate artifact problemhave been proposed by several groups, however all of those proposed todate require rather complicated schemes for acquisition of multipleinterferometric spectra, and none have yet proven satisfactory for highduty cycle imaging. For ophthalmic imaging, the complex conjugateartifact necessitates maintaining very careful positioning of thepatient to avoid overlapping upside-down images (difficult in somepatients), and it precludes imaging tissues any thicker than thesingle-sided imaging depth z_(max) defined next.

Due to spectral sampling considerations, the maximum single-sidedimaging depth z_(max) available in FDOCT is governed by the spectralsampling interval δ_(s)k or δ_(s)λ, according to equation (3) set outbelow:

$\begin{matrix}{z_{\max} = {\frac{\pi}{{2 \cdot \delta_{s}}k} = \frac{\lambda_{0}^{2}}{{4 \cdot \delta_{s}}k}}} & (3)\end{matrix}$

These expressions are given in terms of both wavenumber k (=2π/λ) andwavelength λ (with center wavelength λ₀). In FDOCT systems at 830 nm and1300 nm reported to date, z_(max) has been limited to approximately 4.0mm.

The finite spectral resolution of any real FDOCT system, whethergoverned by the linewidth of a swept laser source in SSOCT, or thegeometric optical performance of the spectrometer convolved with thefinite pixel size of the detector array in SDOCT, gives rise to afalloff in sensitivity with imaging depth that is independent of lightattenuation within the sample. More generally for both SDOCT and SSOCTsystems, if an “effective” detector sampling resolution δ_(r)k or δ_(r)λis defined which accounts for all effects limiting the spectralresolution of the sampled elements, a simpler expression can be derivedfor the falloff to the 6 dB SNR point as illustrated by equation (4) setout below:

$\begin{matrix}{z_{6{dB}} = {\frac{2{\ln (2)}}{\delta_{r}k} = {\frac{\ln (2)}{\pi}\frac{\lambda_{0}^{2}}{\delta_{r}k}}}} & (4)\end{matrix}$

In typical FDOCT systems, the falloff phenomenon exacerbates thealready-limited imaging depth z_(max). In SSOCT systems, the spectrallinewidth is a function of laser dynamics and the detector samplingarchitecture. In SDOCT systems, the spectral sampling interval δ_(s)kand spectral linewidth δ_(r)k are generally a function of spectrometerdesign. In a well-designed spectrometer with Nyquist sampling of theoptics-limited resolution, δ_(r)k≧2δ_(s)k. For the case ofδ_(r)k=2δ_(s)k, equations (3) and (4) can be combined to obtain theuseful rule of thumb for SDOCT systems illustrated by equation (5) setout below:

$\begin{matrix}{\frac{z_{6{dB}}}{z_{\max}} \approx 0.44} & (5)\end{matrix}$

Thus, in practical SDOCT systems, as in most commercial ophthalmic SDOCTsystems, the useful portion of the depth imaging range, defined as the 6dB falloff point, is limited to approximately half of the range given bythe spectral sampling, i.e. approximately 2.0 mm instead of 4.0 mm. Thismay be sufficient for imaging the normal retina, however it may precludeimaging structures above and below normal retina, for example, vitreousfeatures, choroid, and deeply cupped optic nerve heads. It may also beinsufficient for imaging almost any anterior segment structures besidesthe cornea without incurring upside-down artifacts.

The complex conjugate artifacts and falloff of sensitivity with imagingdepth are fundamentally new limitations in FDOCT, which have not yetbeen successfully addressed by technical innovation. These phenomenarepresent significant limitations to the applicability of FDOCTtechniques for ophthalmic diagnostics which require imaging ofstructures deeper than about 2.0 mm.

Three improvements may be combined for overcoming the limitationsdiscussed above and enabling deep imaging FDOCT systems for newapplications of FDOCT where increased depth and removal of mirror-imageartifacts are desirable. Deep-imaging sampling architectures increasez_(max). Modifying the sampled spectral bandwidth such that thebandwidth of the sampled element is less than the sampling intervalreduces the deleterious effects of sensitivity falloff. Addition ofphase information to the acquired spectrum provides informationsufficient to remove complex conjugate artifacts. The combination of thelatter two techniques enables system design to quadruple availableimaging depth without impacting the axial resolution of the imagingsystem. Tailoring the sampling architecture to adjust the maximumimaging depth zmax requires a trade-off between axial resolution andmaximum imaging depth, as shown in FIG. 15, as may be appropriate forthe imaging of target structures. These techniques may be applied toeither SDOCT or SSOCT implementations, and to implementations thatcombine elements of SDOCT and SSOCT. Further, in some cases thesetechniques may be applied dynamically, controlling z_(max),complex-conjugate management, and falloff in situ to manage trade-offsin pixel resolution, region of subject focus, optical power on subject,and imaging speed as may be appropriate for specific objectives duringthe imaging process.

For a traditional volume phase holographic (VPH) grating basedspectrometer design, the imaging depth, as measure in tissue ofrefractive index n, is related to the bandwidth and pixel count of thespectrometer as illustrated in equation (6) set out below:

z _(max)=λ_(c) ²/4nλ _(s)  (6)

Where

-   -   δ=spectrometer bandwidth (nm)=    -   λ_(c)=source center wavelength (nm)    -   p=pixels (detector channels)    -   λ_(s)=spectrometer wavelength spacing=δ/p    -   n=index of refraction

A key spectrometer design decision is to optimize for image resolution,by maximizing available bandwidth δ, or optimize for imaging depth, byminimize sampling interval λ_(s),

With λ_(s)=δ/p equation (6) becomes:

z _(max) =pλ _(c) ²/4nδ  (7)

For fixed p, solving for δ in becomes:

δ=pλ_(c) ²/4nz _(max)  (8)

Alternatively, for fixed δ and solving for λ_(c) in nanometers leads to:

λ_(c)=√(4nδz _(max) /p)  (9)

The determination of the optimum values for λ_(c) and δ are based uponthe requirements for the application, for example, wavelength andimaging depth.

Further definition of the design parameters can be obtained by relatingthe image size to the detector pixel size in order to determine thespectrometer focal length required. Assuming a collimated beam input tothe grating the diffraction limited spot size can be represented by thefollowing expression set out in equation 10:

D=1.22λ_(c)(f/d)  (10)

Where f is the focal length of the spectrometer imaging optics and d istypically the lens aperture diameter which in this case is equivalent tothe spectrometer input collimated beam diameter. Solving for (f/d),

(f/d)=D/1.22λ_(c)  (11)

Therefore, given an exemplary pixel size of 10 um and setting the targetdiffraction limited image spot radius to the 75% of detector pixel sizeyields a spot diameter of 7.5 μm. From equation (11) it can bedetermined that the ratio of the focal to input beam diameter:

(f/d)=3.5  (12)

From expression (10) for a collimated beam of 25 mm in diameter therequired focal length of the spectrometer imaging optics is 89 mm.Conversely, setting the focal length to 100 mm requires a 28 mmcollimated beam input. The determination of which parameter to solve foris based on other physical design constraints of the spectrometer.Further definition of the spectrometer optical design to achieve adiffraction limited spot across a detector array is known and thereforewill not be discussed herein.

Using the grating equation (13):

λ_(c) f=sin θ_(i)+sin θ_(d)  (13)

Where

-   -   λ_(c)=source center wavelength    -   f=spatial frequency of the grating    -   θ_(i)=angle of incidence    -   θ_(d)=angle of diffraction

For standard planar transmission VPH grating designs θ_(i)=θ_(d),Solving for f, equation (13) is reduced to:

f=2 sin θ/λ_(c)  (14)

With the practical upper limit established by:

f=2/λ_(c)  (15)

Since the spectral dispersion of the VPH grating is proportional to thegrating spatial frequency, design optimization is directed toward thespatial frequency. The optical design of the spectrometer is importantin selecting the grating dispersion value. With a spectrometer detectorarray of predetermined physical length and a fixed center wavelength andbandwidth, the dispersion is selected to insure full coverage of thespectral bandwidth across the detector array.

By definition, the dispersion of the grating is the rate of change ofthe angle of diffraction with wavelength for a fixed angle of incidenceor Δθ/Δλ, which from a differentiation of equation (14) yields:

Δθ/Δλ=f/cos θ  (16)

The optical design of the spectrometer and grating dispersion areinterrelated. For a given array length and focal length of the imagingoptics the angle of dispersion can be given as:

φ=2 tan⁻¹(A/2f)  (17)

where f is the focal length of the imaging optics and A is the detectorarray length. From equation (17) the grating dispersion relates thedispersion angle by:

φ=afδ/cos θ=2 tan⁻¹(A/2f)  (18)

where a is the unit conversion from radians/mm to degrees/nm and f isdetermined by the detector pixel size as stated in equation (12).

Using equation (14) the expression can be reduced as follows

(aδ/λ _(e))tan θ=tan⁻¹(A/2f)  (19)

Solving for θ:

θ=tan⁻¹[(λ_(c) /aδ)tan⁻¹(A/2f)]  (20)

From the above equations, the required dispersion angle can becalculated for a given spectrometer layout. The parameters required asinputs to the equations are the detector pixel size which defines therequired focal length, (A) the linear dimension of the detector array,(λ_(c)) the center wavelength of the source and (δ), the bandwidth ofthe source. From the calculated dispersion value, the grating frequencyand grating angle can be calculated resulting in a completecharacterization of the spectrometer design.

Grating-based spectrometer designs as discussed above disperse the lightlinearly as a function of wavelength across the detector array. InFourier Transform Spectroscopy and Fourier Domain Optical CoherenceTomography, the signal of interest is the Fourier transform of thedetected spectrum. The Fourier transform analog to spatial position isspatial frequency, but the detector captures spatial period and thusrequires an additional interpolation step to scale the detected spectrumfrom spatial period to spatial frequency. This resampling is atime-consuming process, and the elimination of such would enable bothfaster processing and more accurate sampling in spatial frequency orwavenumber (k) space.

Additionally, resampling is inadequate to the task of providing aconstant depth scale in the Fourier transformed spatial image. Thechirped sampling (relative to the spatial frequency) yields a chirp indepth per pixel across the image depth. As the imaging window becomesdeeper, this chirp is more deleterious to dispersion compensation and toquantitative measurements across the image depth. It is thereforedesirable to design an FDOCT system sampled linearly in frequency (k,wavenumber).

As indicated, the maximum depth of the SDOCT system is defined by thespatial sampling of the spectrum at the detector—increased wavelength(or wavenumber) sample density allows for sampling of higher frequencyfringes on the spectrum and thus returns signals from deeper depths.This relationship is related by:

$\begin{matrix}{Z_{\max} = \frac{1}{4*\delta \; v_{s}}} & (21)\end{matrix}$

Where δν_(s) is the wavenumber sampling at the detector.

Wavelength and wavenumber are related by:

$\begin{matrix}{\frac{\Delta \; v}{v} = \frac{\Delta \; \lambda}{\lambda}} & (22)\end{matrix}$

Equi-sampling in wavenumber will reduce the burden on computationalresampling, and improve the linearity of the depth scaling in the finalimage. Additionally, application of wavenumber or k-linearization iswell suited to channelized imaging, for example through the use of acomb filter for SDOCT and SSOCT, or through the use of controlledduty-cycle sampling in SSOCT, as discussed below.

Design of k-linearized spectrometers using a prism air-spaced withrespect to a grating has been reviewed elsewhere, for example, inFourier Domain optical coherence tomography with a linear-in-wavenumberspectrometer by Hu et al. However, the use of a prism-air space-gratingconfiguration requires control of extra degrees of freedom, and adds tothe number of glass-air interfaces, potentially reducingmanufacturability and increasing costs. As originally described inConstant-dispersion grism spectrometer for channeled spectra by Traub, aprism-grating (GRISM) structure in intimate contact may be adequate tothe task of creating, in the language of Traub, a constant dispersion(k-linear) spectrograph. Traub, however, does not provide a prescriptionfor practical design of a grism spectrometer that meets the requirementsof FDOCT imaging, including the relationship between required dispersionand degree of linearization required. As shown below, with properspecification of grating spatial frequency, prism index and chromaticdispersion, prism angle, and input angle, a k-linear spectrometer can bedesigned with sufficient linearity to support a frequency-channelizedimplementation with improved sensitivity falloff characteristics.

The exit angle, β, of an isosceles prism is related to the entranceangle, α, the vertex angle, ε, and the index of refraction of the prismas a function of wavelength, n_(p)(λ). Using Snell's law and assumingthe medium surrounding the prism is air, the angle of the light afterrefracting at the first surface of the prism, θ₁, is:

$\begin{matrix}{\theta_{1} = {\sin^{- 1}\left( {\frac{1}{n_{p}(\lambda)}{\sin \left( {\alpha - \frac{ɛ}{2}} \right)}} \right)}} & (23)\end{matrix}$

Following the same logic, the angle after refraction at the secondsurface of the prism, θ₂, is:

$\begin{matrix}{\theta_{2} = {\sin^{- 1}\left( {\frac{n_{p}(\lambda)}{n_{g}(\lambda)}{\sin \left( {\theta_{1} + \frac{ɛ}{2}} \right)}} \right)}} & (24)\end{matrix}$

Where n_(p)(λ) is the wavelength dependent index of refraction of thegrating.

The grating equation is:

$\begin{matrix}{{{\sin \; \alpha} + {\sin \; \beta}} = \frac{{- m}\; \lambda}{d}} & (25)\end{matrix}$

Where α is the angle of incidence onto the grating, β is the exit angleof the grating, m is the diffraction order, λ is the wavelength of theincident light, and d is the groove spacing. Rearranging for the exitangle yields:

$\begin{matrix}{\beta = {\sin^{- 1}\left( {\frac{{- m}\; \lambda}{d} - {\sin \; \alpha}} \right)}} & (26)\end{matrix}$

For a fixed input angle, in the small angle approximation, the angularchange as a function of wavelength is:

$\begin{matrix}{\frac{\beta}{\lambda} = \frac{- m}{d}} & (27)\end{matrix}$

Wavenumber, ν, is:

$\begin{matrix}{v = {\frac{1}{\lambda} = \frac{k}{2\pi}}} & (28)\end{matrix}$

Converting the dispersion equation to wavenumber, ν, yields:

$\begin{matrix}{\frac{\beta}{v} = \frac{- m}{v^{2}d}} & (29)\end{matrix}$

The Sellmeier equation relates the index of refraction n(Aλ) to thewavelength of light using well-characterized, commonly knowncoefficients, B1-3 and C1-3:

$\begin{matrix}{{n(\lambda)} = {1 + \frac{B_{1}\lambda^{2}}{\lambda^{2} - C_{1}} + \frac{B_{2}\lambda^{2}}{\lambda^{2} - C_{2}} + \frac{B_{3}\lambda^{2}}{\lambda^{2} - C_{3}}}} & (30)\end{matrix}$

Equation 30 in terms of wavenumber is:

$\begin{matrix}{{n(v)} = {1 + \frac{B_{1}}{1 - {v^{2}C_{1}}} + \frac{B_{2}}{1 - {v^{2}C_{2}}} + \frac{B_{3}}{1 - {v^{2}C_{3}}}}} & (31)\end{matrix}$

This equation can be used to model the index across the wavelengths orwavenumbers for a given SDOCT wavelength range.

A k-linear GRISM is a combination of a prism and a grating in which thewavenumber dispersion of the prism balances the wavenumber dispersion ofthe grating. This can be tailored to yield approximately constantwavenumber dispersion across the output of the GRISM. One implementationof this design uses an isosceles prism with a flush-mounted VPH gratingas illustrated in FIG. 7 of the present application. Alternative designsthat reverse the order of grating and prism, or that utilize prisms onboth the entrance face and exit face of the gratings may be employedwithout deviating from the invention. Note also that a chirpedholographic grating can be tailored to replicate the transmissionfunction of generally any GRISM. For example, a chirped gratingholographically written will perform as the equivalent GRISM, withoutthe need for mounting a physical GRISM to the grating in thespectrometer. The concept of chirped gratings are discussed in, forexample, U.S. Pat. Nos. 4,834,474 and 7,224,867. Techniques fordesigning a transfer function and preparing a holographic transmissionfilter to provide the targeted transfer function are discussed in, forexample, U.S. Pat. No. 7,519,248. These concepts have not previouslybeen applied to k-linearized spectrometers.

In particular, as illustrated in FIG. 7, P and G are the prism andgrating, respectively. ε is the vertex angle of the prism, α is theangle of incidence onto the prism and β is the deflection angle from theGRISM. As illustrated, d is the width of the detector, for example,20.48 mm. The objective lens L_(obj) represents the optics used to focusthe GRISM output across the detector array. ν_(min) and ν_(max) are thewavenumber range minimum and maximum values, respectively.

FIG. 8 is a graph illustrating grating and prism dispersion inaccordance with some embodiments of the present inventive concept. Thedispersions of grating and prism are additive, such that the utilizationof a prism reduces the dispersive power required of the grating. FIG. 9is a graph illustrating cumulative pixel shift from center pixel due todispersion in accordance with some embodiments of the inventive concept.A non-linearized spectrometer will not support a fullyfrequency-channelized set of spectral elements whereas a k-linearspectrometer can be channelized such that the cumulative offset of theNth frequency channel from the Nth detector pixel is less than onepixel, and preferably less than one-half pixel. FIG. 10 is a graphillustrating GRISM angle dispersion in accordance with some embodimentsof the present inventive concept, demonstrating the linearity withrespect to wavenumber.

Non-ideal spectral sampling in FDOCT systems imposes a depth-dependentfalloff of Signal-to-Noise Ratio (SNR). This falloff is based on thelineshape of the sampled element. For example, if the detected samplingfunction is a square pixel (rect function), then the transform of thesampling function is a sync function, and the shape of the sync functiondefines the falloff window.

Sensitivity falloff is in effect a characteristic of the finitecoherence length of each sampled spectral element. In principle,sampling a comb of single frequencies, for example, a comb of deltafunctions, would completely eliminate sensitivity falloff. This is notachievable in practice. However, a comb convolved with a function, forexample a Gaussian or Lorentzian, whose width is less than the combspacing will demonstrably improve the falloff characteristics; thenarrower the convolving function, or stated alternatively the smallerthe duty cycle of the comb, the greater the positive impact onsensitivity falloff. This effect will be operative for anyimplementation of FDOCT, whether SDOCT or SSOCT, and whether appliedwith a resampled wavelength-sampled spectrum or a k-linear sampledspectrum, though operation in conjunction with a k-linear sampling, suchthat each sampled element records a spectral element of the comb, may bepreferred.

A Fabry-Perot etalon can be used to provide such a comb source. Apractical etalon may be composed of a glass block with 2 partiallyreflecting surfaces. As will be described, the two key attributes of theetalon are the free spectral range (FSR) and the Finesse. The FSRdetermines the sampling interval, which in some embodiments is designedto match the desired sampling interval, for example, the pixel spacingof the k-linear spectrometer or the k-trigger of the SSOCT light source.The FSR is closely related to the optical path length through theetalon. The Finesse sets the spectral width at each output frequency, orthe duty cycle of the etalon transmission function. The Finesse isclosely related to the reflectivity of the interfaces of the etalon.

Light incident etalon, normal to the surface or angled, will either passthrough block or reflect from the block (assuming a lossless etaloninterior). Transmission through the block is defined by:

$\begin{matrix}{T_{e} = \frac{T^{2}}{\left( {1 - R^{2}} \right)\left( \frac{\sinh \; \gamma}{{\cosh \; \gamma} - {\cos \; \delta}} \right)}} & (32)\end{matrix}$

Where T and R are the surface transmission and reflection values,

${\gamma = {\ln \left( \frac{1}{R} \right)}},$

and δ, the phase of the light traveling through the block, is definedby:

$\begin{matrix}{\delta = {\frac{4\pi}{\lambda}{nl}\; \cos \; \theta}} & (33)\end{matrix}$

Where n is the index of refraction of the glass block, λ is thewavelength of the incident light, l is the thickness of the block, andθ□ is the angle of incidence onto the block.

The Free Spectral Range (FSR) of the etalon defines the spacing betweenadjacent transmission peaks and is defined by:

$\begin{matrix}{{FSR} = \frac{\lambda_{0}^{2}}{{2{nl}\; \cos \; \theta} + \lambda_{0}}} & (34)\end{matrix}$

Where λ₀ is the center wavelength of the transmission peak. The FullWidth at Half Maximum (FWHM or Δλ) of each transmission peak is relatedto the finesse,

of the etalon by:

$\begin{matrix}{\mathcal{F} = {\frac{FSR}{\Delta\lambda} = \frac{\pi}{2{\arcsin \left( {1/\sqrt{F}} \right)}}}} & (35)\end{matrix}$

Where F is the coefficient of finesse, which is defined by:

$\begin{matrix}{F = \frac{4R}{1 - R^{2}}} & (36)\end{matrix}$

The thickness of the block and the reflectivity of the surfaces can betailored to provide a comb source for a given wavelength range thatprovides a sub-interval lineshape and a FSR equal to the spectralsampling interval.

The maximum depth of a spectrometer is defined by the frequency spacingat the detector; finer frequency sampling yields a deeper maximum depth.56 nm from 812-868 nm dispersed across 2048 pixels will provide aspectral sampling of 0.027 nm/pixel and a maximum depth of 6.55 mm.Assuming an incident angle of π/8 (22.5⁰) and an index of refraction ofthe etalon glass of 1.55, the FSR and FWHM can be tailored to providesub-pixel FWHM and transmission peak spacing equal to the spectralsampling interval. Assuming a GRISM—based, constant wavenumberdispersion spectrometer is in place, the spectral sampling will beevenly spaced from 1.15×10⁶ in⁻¹ to 1.23×10⁶ m⁻¹. Reflectivity R of 0.24yields a finesse of 1.1, and for a thickness of 10 mm, this yields amean FSR of 0.024 nm and a FWHM of 0.021 nm. Increasing the finesseshortens the FWHM as illustrated below in FIG. 21 and subsequently thefalloff effect as illustrated in FIG. 23, but this also decreases thetotal power output of the source.

For comprehensive FDOCT imaging of the eye by rapidly switching betweenimaging modes designed for imaging different ocular structures along thevisual axis, it would be desirable for the imaging depth (axial field ofview) of each mode to be optimized for the expected length and desiredaxial sampling density of each structure. For example, for imaging ofthe entire anterior segment, the optimal imaging depth is the expectedmaximum anterior segment depth of the anticipated patient population,which may be 6 to 8 millimeters. For imaging of the retina, which isless than about 1.0 mm thick in most locations and contains many closelyspaced layers and structures, it may be preferable for the retinalimaging mode to have a shorter imaging depth and denser axial sampling.

In all FDOCT systems, as has been expressed, there is an inverserelationship between the imaging depth z_(max) and the spectral samplinginterval in wavenumber units δ_(s)k given by:

$\begin{matrix}{z_{\max} = \frac{\pi}{{2 \cdot \delta_{s}}k}} & (37)\end{matrix}$

The total sampled spectral width is given by the spectral samplinginterval δ_(s)k multiplied by the number of spectral samples acquiredper A-scan, typically several thousand, and thus the depth samplingdensity is given by the imaging depth divided by the number of spectralsamples, or some multiple of that number if interpolation is performed.In SDOCT systems, the spectral sampling interval δ_(s)k is typicallyfixed by the spacing of the pixels on the array detector used in thespectrometer and the magnification and spectral dispersion of theinternal optical elements of the spectrometer. In SSOCT systems,however, the spectral sampling interval δ_(s)k is determined by thesweep rate of the light source and/or the electronic sampling rate ofthe analog to digital converter which is recording the SSOCT signal, atleast one of which may be rapidly adjustable electronically or by othermeans. In the case of SSOCT, therefore, it will be desirable to adjustthe spectral sampling interval and thus the imaging depth and depthsampling density (according to the prescription in equation 3) on thefly according to the structure or part of the eye which is being imaged.This imaging depth switching may be coupled to sample and reference armmode switching, such that when switching the sample arm optics andreference arm delay from the anterior segment to the retina, forexample, the imaging depth is also switched to allow for optimal imagingdepth and sampling density of retinal structures. Or, the imaging depthand depth sampling density may be varied within a single operating modeof the sample and reference arm optics, for example to switch betweenshort imaging depth, high spatial sampling density imaging of the corneaand long imaging depth, lower spatial sampling density imaging of theentire anterior segment.

In unmodified SDOCT systems, δ_(r)k is usually limited by the spectralresolution of the spectrometer including the finite spacing of the CCDpixels and diffraction in the spectrometer. In unaltered SSOCT systems,δ_(r)k is typically limited by the instantaneous lineshape of the sweptlaser source, although other factors such as the bandwidth of thedetection electronics may also come into play.

In comprehensive ocular SSOCT systems as described above wherein thespectral sampling interval and depth sampling density are adjusted asper equation 3 according to the structure or part of the eye which isbeing imaged, it is desirable to further implement a comb filter fordecreasing the extent of sensitivity falloff which is also suitablyadjustable to maintain the comb spacing or FSR as the spectral samplinginterval is adjusted. In Fabry-Perot etalons, the FSR is related to thethickness of the etalon, the index of refraction of the material insidethe etalon, and the angle of light incidence upon the etalon. Accordingto some embodiments, one or more of these parameters should be varied insynchrony with changing the spectral sampling interval δ_(s)k in orderto keep the comb filter peaks within their respective spectral samplingintervals. In some embodiments, this may be done by employing a tunableFabry-Perot filter, for example, which utilizes a piezo-electric elementto electronically tune its FSR. Electronic control of the FSR of such afilter may be electronically coupled to the mechanism for changing thespectral sampling interval δ_(s)k, for example by changing thedigitization rate of the analog-to-digital converter.

Note that in such a case the FSR of the comb filter matches the samplingrate of the detector. This is the function of k-triggers commonlydeployed in SSOCT systems to trigger the acquisition of spectralelements. Thus it is conceivable to use the comb filter for a secondaryfunction, to act as the system k-trigger. The converse property does nothold. In particular, a k-trigger is not implemented in current systemsto operate as a comb source generator for the SSOCT system. The proposedcomb filter may be used as a k-trigger in at least two different modes.In a first mode, a small fraction of the transmissive (T) output of thecomb filter is split out of the source path to k-trigger circuitry. Insuch a configuration, the k-trigger implementation is directly analogousto implementations currently used in the art, with the benefit that aseparate device is not required. This mode is fully functional, butcomes at some cost to the power available for imaging.

A second mode is to use the back-reflected (R) light from the filter.The backreflection from a lossless etalon filter is the spectralcomplement to the transmission through the filter, as illustrated inFIG. 24. This backreflection may be used as the k-trigger for an SSOCTsystem. Embodiments illustrating the second mode including a sweptsource followed an optical isolator, an etalon filter, and an opticalcirculator, will be discussed further below with respect to FIG. 19. Thebackreflected output from the etalon is directed to k-trigger circuitryand applied to trigger spectral sampling of a balanced heterodynedetector. The balanced detector sees interference signature both fromthe detector port of the coupler, and the shunt port of the opticalcirculator.

To resolve the complex conjugate artifact, several academic groups havepointed out that a second spectral interferogram may be obtained withthe phase offset φ shifted in phase by π/2. Combining the real andimaginary parts yields the complex interferometric signal {circumflexover (D)}_(i)[k_(m)]={circumflex over (D)}_(i) ⁰[k_(m)]+j{circumflexover (D)}_(i) ⁹⁰[k_(m)], the Fourier transform of which reveals anA-scan with the position of the sample arm reflector unambiguouslydetermined. A method to obtain the complex signal using only two phasestepped scans has been demonstrated, but completely artifact-free tissueimaging has only been demonstrated using a 5 step algorithm in which theadditional phase steps were necessary to compensate for phase errors.

For an SDOCT system embodiments for complex conjugate removal (CCR) maybe via sinusioidal phase modulation as discussed in, for example,commonly assigned U.S. Pat. No. 7,742,174, the disclosure of which hasbeen incorporated herein above. In particular, the system discussed inaccordance with some embodiments of U.S. Pat. No. 7,742,174 isillustrated in FIG. 11. Referring to FIG. 11, the optical coherencetomography (OCT) system 1100 includes a piezoelectric transducer (PZT)element. As illustrated in FIG. 11, the system 1100 further includes alight source 1110, a detector 1120, a fiber coupler 1130, a referencedelay 1140, a piezo-mirror combination 1190, a beam steering unit 1160,a sample arm 1150 and a sample 1170. The light source 1110 may include abroadband light source and the detector 1120 includes a spectrometerilluminating a multichannel detector, such as a linear charge-coupleddevice (CCD) array. A piezo-mirror combination 1190 is located in thereference arm 1140 of the interferometer, which may include a mirror1191 and a piezoelectric element 1192 as illustrated therein.

As discussed in U.S. Pat. No. 7,742,174, phase modulation (linearcontinuous phase modulation 1101A or sinusoidal continuous phasemodulation 1101B) involves placement of a path length modulation ineither the sample or reference arm of an SDOCT system which varies thedifferential path length between the arms with amplitude and phase givenin the text preceding equation (14) in U.S. Pat. No. 7,742,174, at arate corresponding to π/4 radians of phase modulation per A-scanintegration time of the spectrometer. Then, each set of four sequentialA-scan acquisitions are combined according to equation (14) of U.S. Pat.No. 7,742,174 in order to generate an A-scan with total depth equal to2*z_(max) as defined above. If the amplitude, phase and frequency of themodulation are set as specified in U.S. Pat. No. 7,742,174, then theresulting A-scan should theoretically be completely free of DC,autocorrelation, and complex conjugate artifacts.

However, slight deviations from perfection in achieving these parameterssuch as will be experienced in any real physical implementation ofsinusoidal phase modulation may lead to a degradation of performancecompared to the ideal result in the form of incomplete complex conjugateartifact suppression. Thus, an additional step of applying quadratureprojection processing as discussed with respect to FIG. 2 of U.S. PatentApplication Publication No. 2008/0170219 may be applied to improve thecomplex conjugate artifact rejection, at the cost of a small amount ofreduced signal to noise ratio. Quadrature projection processing is analgorithmic step which does not require any hardware modification andwhich reduces the complex conjugate artifact from imperfectly phasemodulated SDOCT data by forcing the real and imaginary parts of therecorded A-scan signal to be orthogonal.

For an SSOCT system, some embodiments implement complex conjugateremoval (CCR) using the heterodyne CCR method as discussed in commonlyassigned U.S. Pat. No. 7,336,366, which involves introducing a frequencyshift between the sample and reference arm light and thus shifting thecarrier frequency of the image-bearing signal away from DC, about whichthe complex conjugate artifact is centered. With the addition of thisfrequency shift, the A-scan free of complex conjugate artifact is foundfrom the Fourier transform of the detected signal, centered at thefrequency shift value. If an A/D converter is used which has much higherbandwidth than the SSOCT signal itself, then the frequency shift valuecan be set to be many times the frequency encoding the z_(max) value ofthe A-scan, thus the complex conjugate artifact will be located far infrequency space away from the A-scan data. If a very high sweep speed isused, however, such that the SSOCT signal already occupies a substantialfraction of the A/D converter bandwidth, then the complex conjugateartifact may only be shifted to the borders of the depth-doubled A-scan.This method of heterodyne CCR is consistent and will not interfere withthe embodiments described above for filtering to improve sensitivityfalloff and sampling to adjust maximum single-sided imaging depth.

Some embodiments of the present inventive concept are directed tocomprehensive volumetric imaging of all ocular structures along thevisual axis using Fourier-domain optical coherence tomography (FDOCT).Current-generation FDOCT systems, including spectral-domain (SDOCT) andswept-source (SSOCT) implementations, are in routine clinical use fordiagnosis of retinal pathologies. FDOCT systems have also been appliedfor imaging of the anterior segment of the eye. Existing optical designsfor scanning the anterior segment and retina are illustrated in FIGS. 1through 3 of the present application. FDOCT is useful for examination ofthe anterior segment of the eye, for diagnosis of corneal, iris, andlens pathologies as well as for quantitative biometry of the anteriorsegment including measurements of corneal refractive power, cornealthickness, anterior chamber depth, lens optical power, and lensthickness. These parameters resulting from anterior segment biometry,with the addition of eye length measurement, are needed for calculationof intraocular lens implant power for cataract surgery. Current methodsfor evaluation of these parameters are limited to measurement along asingle axis, and thus provide only central values for these parameterswhich may not accurately account for off-axis variations andaberrations. With the ability to rapidly acquire densely sampled 2Dimages and 3D volumes of information, FDOCT offers the potential toperform substantially improved characterization of the refractiveproperties of the entire eye, if calibrated and correlated volumetricimages of the anterior segment, lens, and retina could be acquiredeither simultaneously or in rapid succession in the same patient.

Current-generation FDOCT instruments, however, are not capable ofimaging with sufficient depth field of view to record data from all ofthese structures with the same instrument without time-consuminginterchange of optics and of the reference arm length. Thus, there is aneed for FDOCT system designs capable of either simultaneous imaging ofthe anterior segment, lens, and retina or of rapidly switching betweensuch modes during a rapid acquisition sequence which preserves theirrelative displacements in order to perform comprehensive volumetricimaging of all ocular structures along the visual axis. Such switchingshould preferably be rapid, on the time scale of a few A-scansacquisition time, i.e. a few milliseconds, and should allow for themaximum possible re-use of optics and mechanics in both modes to reducetotal system cost and complexity.

Applying the techniques described in this inventive concept, adynamically adjustable extended depth imaging system may be applied toophthalmic imaging for targeted imaging of any region of the eye withoptimized depth field of view and image resolution. FIGS. 12 and 13A-13Eillustrate a series of imaging windows 1255 and 1355-1355″″ that may beapplied for a select variety of imaging circumstances, for example,vitreoretinal surgery, cataract surgery, cornea and anterior chambersurgery and the like. As illustrated in FIGS. 12 and 13A-13E, the seriesof windows may have a variety of sizes, shapes and locations inaccordance with embodiments discussed herein.

In particular, FIG. 12 illustrates using a normal imaging depth window1255 to image the whole eye. As illustrated in FIG. 12, to image thewhole eye using window 1255, six depths would have to be taken to obtainimages of the whole eye. With each depth, a focal adjustment andreference arm adjustment is made.

Referring now to FIG. 13A, as illustrated therein, using phasemodulation and complex conjugate techniques discussed below, a windowhaving a double depth 1355 may be used, which can decrease the number ofsteps from six in FIG. 12 to three steps in FIG. 13A. Each step alsorequires focal and reference arm adjustments, which can be fine tunedwith continuous adjustments to the reference arm.

Referring now to FIG. 13B, as illustrated therein, using an extendeddepth window 1355′ in accordance with some embodiments also allows thenumber of steps to be reduced from six steps in FIG. 12 to three in FIG.13B. The extended depth windows 1355′ are provided without the use ofcomplex conjugate techniques. The front and reference arm optics areequivalent and, therefore, does not require phase modulation.

Referring now to FIG. 13C, using both techniques discussed above withrespect to FIGS. 13A and 13B, i.e., a double depth and extended depthwindow 1355″, the number of steps can be further reduced to two.

As illustrated in FIGS. 13D and 13E, the number of steps can be furtherreduced to one step. In particular, as illustrated in FIG. 13D, a singledouble depth window 1355″′ can be used to image the whole eye.Alternatively, as illustrated in FIG. 13E, a single extended depthwindow 1355″″ can be used to image the whole eye.

FIGS. 12-13E are intended to provide example techniques of how a wholeeye can be imaged using various techniques. It will be understood thatother techniques, number of steps and the like can be used withoutdeparting from the scope of the present inventive concept.

Referring now to FIG. 14, a block diagram illustrating an extended depthFDOCT system in accordance with some embodiments of the presentinventive concept will be discussed. As illustrated in FIG. 14, thesystem includes a source 1400, a reference arm 1410 and a sample arm1440 coupled to each other by a beamsplitter 1420. As furtherillustrated in FIG. 14, the beamsplitter 1420 is also coupled to afrequency sampled detection module 1431 over a detection path 1406 thatmay be provided by an optical fiber.

As further illustrated in FIG. 14, the source 1400 is coupled to thebeamsplitter 1420 by a source path 1405. The source 1400 may be, forexample, a broadband comb source. The reference arm 1410 is coupled tothe beamsplitter 1420 over a reference arm path 1407. Similarly, thesample arm 1440 is coupled to the beamsplitter 1420 over the sample armpath 1408. The source path 1405, the reference arm path 1407 and thesample arm path 1408 may all be provided by optical fiber.

In some embodiments, the reference arm 1410 may be a phase modulatedreference arm or a frequency-shifted reference arm as illustrated inFIG. 14, although embodiments of the present inventive concept are notlimited to this configuration. Furthermore, the sample arm 1440 mayinclude scanning delivery optics and variable optics 1460. Alsoillustrated in FIG. 14 are the reference plane 1450 and a representationof a depth doubled imaging window 1470 in accordance with someembodiments of the present inventive concept.

FIG. 15 is a graph illustrating depth and resolution vs. spectrometerbandwidth and samples for an extended depth FDOCT system in accordancewith some embodiments discussed herein. FIG. 16 is a graph illustratingimage depth and sampling free spectral range vs. spectrometer bandwidthfor an extended depth FDOCT system in accordance with some embodimentsof the present inventive concept. FIG. 15 illustrates the relationshipbetween resolution and total imaging bandwidth, given single sidedimaging at 2048 samples and 4096 samples, and complex conjugate resolvedimaging at 4096 samples. As the bandwidth is constrained to increaseimage depth, resolution suffers. FIG. 16 illustrates the same bandwidthand sampling dependence of image depth, as well as the effective freespectral range associated with k-linearized sampling.

Embodiments of the present inventive concept directed to spectral domainOCT (SDOCT) will now be discussed. It will be understood that both SDOCTand SSOCT implementations will be discussed in detail herein. Theselection of SDOCT or SSOCT is a function of desired imaging wavelength,availability of sources, and tradeoffs between key attributes, such asimaging speed and resolution. Implementations have been shown in the artthat combine elements of SDOCT and SSOCT, and such implementations maybenefit from application of the present inventive concept.

Referring again to FIG. 14, an SDOCT system in accordance withembodiments discussed herein includes a broadband optical source 1400, asource path 1405, a beam splitter/combiner 1420, a reference path 1407,a reference reflector 1410, a sample path 1408 with a scanning systemand focal optics 1440/1460 configured to appropriately to imagestructures of the sample, such as the cornea, anterior chamber, iris,lens, posterior chamber, and retina of the eye, a detector path 1407,and a spectrographic detection system 1431.

In some embodiments, the SDOCT system is designed to image structures ofthe eye in the 800 nm to 900 nm wavelength range. The system may bedesigned to have a single-sided imaging depth (as measured in air) ofabout 7.0 mm, suitable for imaging the crystalline lens of the eye, anda complex-conjugate resolved imaging depth of about 14.0 mm, suitablefor full range imaging of anterior of the eye, from corneal apex throughthe crystalline lens. Through translation of the reference arm 1407 andchange in scanning and focal attributes of sample arm optics, the systemis capable of imaging the entire optical structure of the eye in threesteps.

In some embodiments, the broadband optical source 1400 is asuperluminescent diode with a bandwidth of between about 40 nm and about80 nm. The bandwidth of the source may be selected for axial resolution,but the useful bandwidth may be constrained by the total bandwidth ofthe detector. In some embodiments, the spectral characteristics of thesource are such that the spectral power density at the edges of thespectrometer are attenuated at least about 6 dB from the peak powerdensity, and may be about 10 dB. If the optical power at the edges ofthe spectrometer is too high, the image may exhibit ringing aroundbright features; numerical windowing of the acquired spectrum willreduce this artifact. The parameters of the numerical windowing may beselected to reduce the ringing by smoothly attenuating the signal tomeet the stated conditions. For example, a cosine-squared window may beapplied to the data immediately prior to the Fourier transform, or araised Gaussian function may be applied (e^(−x̂4)).

Although embodiments are discussed herein as having a superluminescentdiode for the broadband optical source 1440, embodiments of the presentinventive concept are not limited to this configuration. However, thesuperluminescent diode may be the most cost effective in thisapplication, where ultra-wide bandwidth may not be required.

In some embodiments, the paths may be combined using single-mode opticalfiber, such as Corning HI780. A fiber optic coupler may be used as thebeam splitter/combiner 1420. The splitting ratio of the coupler can bechosen to optimize power to the sample and signal-to-noise ratio of thedetection system. In some embodiments, the splitter 1420 may have a80/20 split ratio, with 20% of the source light directed to the sampleand 80% directed to the reference arm.

The reference path directs light from the coupler to an opticalreflector. The path length of the reference arm may be designed to matchthe path length to the region of interest for the sample under test. Insome embodiments, the reference arm 1407 has a translation capability toadjust to varying regions for a sample under test, which may beparticularly important for imaging at multiple depths within one sample,such as an eye. The reference arm 1407 may be continuously translated,translated in steps through switches to predetermined path lengths, or acombination of the two without departing from the scope of the presentinventive concept. Generally, the reference arm may be finely adjustableto a precision of at least about 100 μm to accurately position thesample within the FDOCT imaging window 1470.

The sample arm 1408 includes scanning optics, preferably scannersconfigured to scan a beam to any position within a field of view;scanning may be continuous, as with galvonometric scanners, ordiscontinuous, using, for example projecting a beam onto a spinningdiffractive structure without departing from the scope of embodimentsdiscussed herein. The optics used to deliver the scanned beam to thesubject are discussed in, for example, U.S. Patent Publication No.2008/0106696 incorporated by reference above, for imaging of theanterior structures of the eye, nominally telecentric scanning focusedonto anterior structures, or scanning design to pivot in the pupil ofthe eye for scanning an imaging of posterior structures.

The spectrographic system images the output of the dispersedinterference signal onto a CCD (e.g., Atmel EM2, DALSA Spyder) or CMOS(e.g. Basler Sprint) camera, as is well known in the art. For extendeddepth imaging with 7 mm single-sided imaging depth, a source withcentral wavelength of 840 nm and a FWHM bandwidth of 65 nm imaged onto a4096 element array with 14 μm pixel width may be used. As outlined inthe Table of FIG. 17, the edged-to-edge bandwidth of the array is 103nm, and the source decays to 6 dB of peak power at the edge of thearray. The frequency spacing of central pixels is 10.7 GHz. In atraditional spectrometer that utilizes a volume phase holographtransmission grating, there may be significant frequency chirp from theblue edge to the red edge, leading to the need for resampling discussedearlier.

In some embodiments of the present inventive concept, the spectrometerwill be of a constant-dispersion, or k-linearized type as illustrated inFIGS. 18A through 18C, k-linear spectrometer 1832. As illustrated inFIGS. 18A-18C, the spectrometer includes a comb source 1801, a referencearm 1810 and a sample arm 1890 coupled to each other by acoupler/beamsplitter 1820. As further illustrated in FIGS. 18A-18C, thebeamsplitter 1820 is also coupled to the K-linear Spectrometer over adetection path 1806 that may be provided by an optical fiber. As furtherillustrated in FIG. 18, the source 1801 is coupled to the beamsplitter1820 by a source path 1805; the reference arm 1810 is coupled to thebeamsplitter 1820 over a reference arm path 1807. Similarly, the samplearm 1890 is coupled to the beamsplitter 1820 over the sample arm path1808. The source path 1805, the reference arm path 1807 and the samplearm path 1808 may all be provided by optical fiber. It will beunderstood that this may be performed by replacing the VPH grating witha GRISM—a grating-prism pair discussed above with respect to FIGS.7-10—or a chirped grating replicating a GRISM as discussed above. Insome embodiments, the prism is a high index glass (Schott P-SF68,n=2.0), with a vertex angle of π/8 radians (FIG. 7). The prism is inoptical contact with the grating. The grating is a low-spatial frequencygrating (400 lines/mm), sandwiched between faces of Schott B-270(n=1.52). The prism angle of incidence α is 22.5 degrees. A high indexprism is typically necessary in order that the total internal reflectioncondition of the grating is reduced or possibly avoided. An air-spacedprism-grating combination may be used to provide additional designfunctionality, but is not necessary in all cases. The collimated beaminput to the prism may be 25 mm in diameter. The dispersed output fromthe grating couples to a 100 mm focal system, yielding a <10 micrometerspot size on the pixels across the array. The Nth frequency channel mapsto the Nth pixel to within 50% of the pixel width across the array.

As illustrated in FIG. 18A, the comb source 1801 may include a broadbandsource 1800 and a periodic filter 1803 connected through a path 1802. Insome embodiments, the spectrum may be channelized to the spectrometerusing the periodic optical filter 1803 illustrated in FIG. 18A. In someembodiments, the filter 1803 may be a fabry-Perot etalon (discussedabove) illustrated in FIG. 20, which will be discussed further below. Insome embodiments, the filter 1803 may be an AR coated glass block ofindex 1.55 with FSR of 10.7 GHz and Finesse of two. Operating at angleof π/8 degrees to normal to avoid backreflections into the diode, thethickness of the block is 9.79 mm. To achieve a finesse of 2, thereflectivity of the AR coatings must be 41%. For a finesse of 8,reflectivity is 92.7%, further improving sensitivity falloff, but at thecost of required source power. As the linearity of the spectromter willbe calibrated, precision of the central frequency of the etalon as areference point is not a primary concern. Athermalization may berequired not so much to control shifts in the channelized spectrum, butto control changes to FSR. Athermalization techniques are known in theart; the degree of athermalization required is to keep the FSR constantto within 25%. An alternative to an athermalized glass block is to use apiezo controlled cavity; the cavity spacing would increase to 15.2 mmfor an air index n=1.

The combination of the k-linear spectrometer and the filtered sourcebandwidth yields a (single-sided) deep imaging SDOCT system withsuperior sensitivity falloff characteristics. The addition of phasemodulation to the reference arm as discussed in U.S. Pat. No. 7,742,174or U.S. Patent Publication No. 2008/0002183. In some embodiments, apiezo-driven retrorefelector 1811 as illustrated in FIG. 18C modulatesthe phase of the reference arm from its nominal position. In principle,the phase of the reference arm can be modulated in steps of π/4 foracquisition multiple phase-stepped acquisitions at a specific A-scanlocation.

In practice, to continuously modulate the scan; the phase informationcan be determined by integrating over the π/4 steps using an integratingbuckets approach. Note that it may not be necessary for the phase stepsto be π/4; π/3, for example, works as well. The optimal number of stepsis a function of the level of isolation between the real and the mirrorimage, and the phase stability of the subject. To the latter point,rapid image acquisition may be preferred. In some embodiments, a CMOS orCCD camera with acquisition speeds of at least 70 kHz are desired. In afour phase-step acquisition, a single A-scan is acquired at 17 kHz,which is suitably fast for real-time display of full range crosssectional images. As cameras are now available at 140 kHz, a target fullrange line rate of 34 kHz (1000 line frame rate of 34 Hz) is practical.

Note as well that it may not be necessary that that the scanning mirrorsremain fixed at a specific A-scan location. Phase modulation andacquisition of sequential A-scans is acceptable so long as the A-scansare optically oversampled at a similar ratio as implied in theper-A-scan acquisition scenario. Thus sinusoidally scanning over itradians at each A-scan and acquiring four samples is functionallyequivalent to linearly modulating at a rate of π radians over foursequential 4× oversampled A-scans.

If the amplitude, phase and frequency of the modulation are set asspecified in U.S. Pat. No. 7,742,174, then the resulting A-scan shouldtheoretically be completely free of DC, autocorrelation, and complexconjugate artifacts. However, slight deviations from perfection inachieving these parameters may be experienced in any real physicalimplementation of sinusoidal phase modulation and may lead to adegradation of performance compared to the ideal result in the form ofincomplete complex conjugate artifact suppression. Thus, an additionalstep of applying quadrature projection processing according to FIG. 2 ofU.S. Patent Application Publication No. 2008/0170219 may be applied toimprove the complex conjugate artifact rejection, at the cost of a smallamount of reduced signal to noise ratio. Quadrature projectionprocessing is an algorithmic step which does not require any hardwaremodification and which reduces the complex conjugate artifact fromimperfectly phase modulated SDOCT data by forcing the real and imaginaryparts of the recorded A-scan signal to be orthogonal.

FIG. 25 is a block diagram illustrating data flow in some embodiments ofSDOCT imaging systems in accordance with embodiments discussed herein.As illustrated, the prime bottleneck to stream-to-disk acquisition isnot the PCI Bus or motherboard memory bus but the hard drive bus, whichis typically limited to 300 MB/s per bus for a SATA drive. FIG. 26illustrates a CCR control timing diagram. As illustrated therein, everyfourth line clock is phase locked to the mirror drive and as such thepiezo sync output is locked to the line output.

Referring now to FIG. 19, an SSOCT system designed for comprehensiveocular imaging according to some embodiments of the present inventiveconcept will be discussed. In some embodiments, complex conjugateremoval (CCR) is the so-called “heterodyne” CCR method, which involvesintroducing a frequency shift between the sample and reference arm lightand thus shifting the carrier frequency of the image-bearing signal awayfrom DC, about which the complex conjugate artifact is centered asdiscussed in U.S. Pat. No. 7,336,366. With the addition of thisfrequency shift, the A-scan free of complex conjugate artifact is foundfrom the Fourier transform of the detected signal, centered at thefrequency shift value. If an A/D converter is used which has much higherbandwidth than the SSOCT signal itself, then the frequency shift valuecan be set to be many times the frequency encoding the z_(max) value ofthe A-scan, thus the complex conjugate artifact will be located far infrequency space away from the A-scan data. If a very high sweep speed isemployed, however, such that the SSOCT signal already occupies asubstantial fraction of the A/D converter bandwidth, then the complexconjugate artifact may only be shifted to the borders of thedepth-doubled A-scan. This method of heterodyne CCR is consistent andwill not interfere with the embodiments described above for switchingbetween sample and reference arm imaging modes, switching SSOCT imagingdepth, and switching of the comb filter FSR spacing to remain consistentwith the spectral sampling interval.

As illustrated in FIG. 19, the SSOCT system includes a swept comb source1995, a circulator 1999, a beamsplitter 1920, a triggered balancedheterodyne detector 1933, a frequency-shifted reference arm 1912 andscanning delivery optics 1940 in the sample arm. As further illustratedin FIG. 19, a fabry-Perot etalon (discussed above) 1997 and a sweptsource 1996 can be used to provide a swept comb source. A practicaletalon may be composed of a glass block with 2 partially reflectingsurfaces. As discussed above, the two key attributes of the etalon arethe free spectral range (FSR) and the Finesse. The FSR determines thesampling interval, which in some embodiments is designed to match thedesired sampling interval, for example, the pixel spacing of thek-linear spectrometer or the k-trigger 1998 of the light source. The FSRis closely related to the optical path length through the etalon, whichmay be angle tuned according to equation 34. The Finesse sets thespectral width at each output frequency, or the duty cycle of the etalontransmission function. The Finesse is closely related to thereflectivity of the interfaces of the etalon. As further illustrated inFIG. 19, it is further advisable to use an optical isolator orcirculator 1999 after the filter and before coupler 1920, as signalreturned from sample and reference arm will experience a complementaryinteraction with etalon, and multi-path interference may degrade imagequality.

FIG. 20 is a detailed block diagram of the periodic filter 1997 of FIG.19. Also illustrated in FIG. 20 are graphs of reflected (R) andtransmitted (T) power that is output from the filter. FIG. 21 furtherillustrates a graph depicting the output of the periodic filter of FIG.20. FIG. 22 is a graph illustrating an effective duty cycle of theperiodic optical filter of FIG. 20. FIG. 23 is a graph illustrating SNRfalloff as a function of pixel fill factor (duty cycle). As illustratedtherein, as the fill factor decreases from 100% down to 50% (Finesse=2),for the 7 mm single-side imaging system discussed herein, the 3 dBfalloff depth increases from about 1.34 mm to nearly 2.68 mm. Decreasingfurther to a 15% fill factor pushes the 3 dB depth beyond the maximumdepth. Coupled with CCR, this technique could increase the total imagingrange with SNR loss to a full 14 mm range with 1.8 dB SNR loss at theedges for Finesse=6. Finally, FIG. 24 is a graph comparing reflected andtransmitted power of the optical filter of FIG. 20.

Some embodiments for a comprehensive ocular imaging system using sweptsource (SSOCT) design have a z_(max)=7 μm, thus the imaging depthcapability of this system after complex conjugate removal is 14 mmoptical path length. As illustrated in FIG. 19, for a swept sourceimplementation, the light source may be a swept source laser 1996 havinga center wavelength near 1060 nm, an instantaneous coherence length(before filtering) of 5 mm, and a full-scanning optical bandwidth ofapproximately 100 nm. Light from the laser is directed into a 50:50single mode coupler 1920 and then into sample and reference arms.

As in the SDOCT implementation, the reference path directs light fromthe coupler to an optical reflector 1912 that is designed to match thepath length to the region of interest for the sample under test.Positioning capabilities of the SSOCT reference arm are the same as forthe SDOCT reference arm. However, in some embodiments, rather than thephase modulator of the SDOCT configuration, the SSOCT configurationpossesses an acousto-optic modulator (AOM) operating at 250 MHz acousticfrequency for heterodyne complex conjugate artifact removal. The samplearm may also possess an AOM operating at 250 MHZ plus a differentialfrequency, as discussed in U.S. Pat. No. 7,336,366.

Light returning from the sample and reference arms is recombined in the2×2 coupler and detected by a 500 MHz bandwidth optical photoreceiver.A/D conversion is performed with 12 bit resolution at 500 MHz samplingrate in order to obtain 2*z_(max)=14 mm optical path length.

Previously demonstrated implementations of heterodyne complex-conjugateremoval in SSOCT systems utilized a pair of phase modulators (eitheracousto-optic or electro-optic) arranged to give a net difference phasemodulation frequency on the order of hundreds of kHz to tens of MHz.This was done with either one modulator placed in each of the referenceand sample arms, or two modulators arranged in series in a single arm.With source sweep frequencies of less than about 20 kHz, thisarrangement gives a sufficiently high heterodyne modulation frequency toallow for good separation of the complex-resolved A-scan signal awayfrom DC. With an increased sweep rate of approximately 100 kHz, a singleacousto-optic or electro-optic modulator operating at approximately350-500 MHz modulation frequency may be placed in the reference arm, asillustrated in FIG. 19. If the photoreceiver and A/D conversioncircuitry have a bandwidth of 700-1000 MHz, then the frequencymodulation will place the zero path length position of the A-scan nearthe middle of the detection bandwidth, thus effectively resolving thecomplex conjugate artifact for these rapid scan rates.

The same periodic filter structure described for the SDOCT system isapplied to the SSOCT to increase the instantaneous coherence length ofthe source (by reducing the sampled linewidth). A variable lengthpiezo-driven etalon may be used in order that the frequency spacing ofthe output peaks may be changed to change the single-sided depth of theimage. At 10.71 GHz, a 7 mm single-sided window imaging window may beachieved. The number of samples acquired determines the wavelength rangeutilized, and thus enables a tradeoff between resolution and acquisitionspeed. At 2048 samples, the sampled wavelength range will be 82 nm, andthe resolution will be approximately 10 μm. The reflective port of theperiodic filter acts directly as the k-trigger for sampling theinterference signature. As the etalon FSR is modified, for example from10.7 GHz to 5.35 GHz, the single-sided imaging depth is increased from 7mm to 14 mm. The k-trigger automatically tracks. This capability tochange imaging depth is an important attribute of this SSOCTarchitecture, allowing an imaging system to rapidly change depth ofimaging field as the situation requires.

For SDOCT, one can imagine a hardware switchable spectrometer whereinthe sampling interval is modified. A simple approach to reduce imagedepth is to process every second pixel on an array. In some embodiments,a spectrometer can be constructed to double the imaging depth.

Further embodiments of the present inventive concept will now bediscussed with respect to FIGS. 27 through 41. In particular, variousdetails optics will be discussed for comprehensive ocular FDOCT. Asdiscussed above, some embodiments discussed herein are related tocomprehensive volumetric imaging of all ocular structures along thevisual axis using FDOCT. As further discussed above, current-generationFDOCT systems, including spectral-domain (SDOCT) and swept-source(SSOCT) implementations, are in routine clinical use for diagnosis ofretinal pathologies. FDOCT systems have also been applied for imaging ofthe anterior segment of the eye. As used herein, the “anterior segmentof the eye” refers to the region of the eye from the posterior surfaceof the crystalline lens to the apex of the cornea. Thus, the “anteriorsegment of the eye” refers to all ocular structures located anterior tothe vitreous humor (including cornea, aqueous humor, iris, ciliary body,and crystalline lens). As used herein, the “posterior segment of theeye” refers to the vitreous from the posterior surface of thecrystalline lens up to and including the retina, choroid, and opticnerve. Thus, the “posterior segment of the eye” refers to the internalocular structures which are located posterior to the anterior segment,including the vitreous humor, retina, and choroid. Conventional opticaldesigns for scanning the anterior segment and retina are illustrated inFIGS. 27A through 27C.

In particular, FIG. 27A illustrates a system for imaging the anteriorsegment of the eye 2760. As illustrated, the system includes acollimator 2700, a two-dimensional galvanometer scanner 2710, and asingle scan lens 2720 in a telecentric configuration. As used herein,“telecentric” or “telecentricity” refers to scanning a beam such thatthe rays of the beam are parallel to an optical axis of the system. Asfurther illustrated in FIG. 27, the single scan lens 2720 is coupled tothe sample arm fiber tip 2730. FIG. 27B illustrates a conventionalsystem for retinal scanning with an iris pivot including an additionallens 2701. FIG. 27C illustrates a conventional telecentric anteriorsegment system including a corneal adapter. As used herein, the“collimated” refers to a non-diverging, spatially coherent beam. Inother words, “collimated” refers to a parallel beam (i.e., neitherconverging nor diverging).

FDOCT is useful for examination of the anterior segment of the eye,diagnosis of corneal, iris, and lens pathologies as well as forquantitative biometry of the anterior segment including measurements ofcorneal refractive power, corneal thickness, anterior chamber depth,lens optical power, and lens thickness. These parameters resulting fromanterior segment biometry, with the addition of eye length measurement,are needed for calculation of intraocular lens implant power forcataract surgery. Current methods for evaluation of these parameters arelimited to measurement along a single axis, and thus provide onlycentral values for these parameters which may not accurately account foroff-axis variations and aberrations. With the ability to rapidly acquiredensely sampled 2D images and 3D volumes of information, FDOCT offersthe potential to perform substantially improved characterization of therefractive properties of the entire eye, if calibrated, and correlatedvolumetric images of the anterior segment, lens, and retina could beacquired either simultaneously or in rapid succession in the samepatient. Current-generation FDOCT instruments, however, are not capableof imaging with sufficient depth field of view to record data from allof these structures with the same instrument without time-consuminginterchange of optics and of the reference arm length.

Accordingly, as will be discussed below with respect to FIG. 28 through41, some embodiments of the present inventive concept provide an FDOCTsystem capable of simultaneous imaging of the anterior segment, lens,and retina or of rapidly switching between such modes during a rapidacquisition sequence, which preserves their relative displacements inorder to perform comprehensive volumetric imaging of all ocularstructures along the visual axis. In some embodiments, switching betweenmodes is rapid, i.e., on the time scale of a few A-scans acquisitiontime, for example, a few milliseconds, and should allow for the maximumpossible re-use of optics and mechanics in both modes to reduce totalsystem cost and complexity.

Some embodiments of the present inventive concept configured to performrapid switching between imaging in the anterior segment (including thecornea, acqueous humor, iris, ciliary body, and lens) and retina inFDOCT systems will be discussed. As a preliminary note, embodimentsdiscussed with respect to FIGS. 28A-30B represent optical designs ofpatient ocular scanners which are configured to be placed in the samplearm of FDOCT, SDOCT, or SSOCT systems. In the illustrated embodiments,all lenses are assumed for clarity to have focal length f and to beplaced distances apart as indicated. However one of skill in the art ofoptical design would understand that other focal lengths of some lensescould also be used, with accompanying effects of magnification orde-magnification of the scan patterns and resulting spot sizes, withoutdeparting from the scope of embodiments discussed herein. Furthermore,in the figures, galvanometer scanners are illustrated in a compactformat which is known to those familiar with the art of scanning systemdesign. Two-dimensional galvanometers are shown as crossed dashed lines,denoting that they scan in two dimensions. This may be accomplished witha single mirror which is capable of being pivoted in two orthogonaldirections, or with two separate one-dimensional galvonometers, eachcapable of being pivoted in orthogonal directions which are either placein close proximity (i.e., less than about 10 mm apart), or else havingan optical sub-system placed between them whose purpose it is to relayan image of one mirror onto the other one. In addition, thegalvanometers are also drawn in an unfolded manner so that the lightbeams incident upon them in their undeviated or “central” position areshown as passing straight through rather than being reflected, and lightbeams which are deflected in either one direction from this centralposition are shown either above or below the undeviated beams.

Referring first to FIG. 28A, a system for imaging of the anteriorsegment includes a collimator 2800, a two-dimensional galvanometerscanner 2810, and a single scan lens 2820 in a telecentricconfiguration. As further illustrated the scan lens is connected to thesample arm fiber tip. In the configuration illustrated in FIG. 28A, twodimensional and three dimensional imaging of the entire anterior segmentmay be performed. To switch the system of FIG. 28A to a system forretinal imaging, a single additional lens 2822 as illustrated in FIG.28B, also with focal length f, is rapidly translated into the opticalpath either immediately proximal or immediately distal to thecollimating lens. This lens changes the sample arm beam from collimatedto focusing on the 2D galvanometer scanner. As used herein, “focusing”refers to refracting or steering a beam of light to converge at thefocal position of the focusing element. In other words, “focusing”refers to a beam in which the rays are converging to a common focalpoint. With the scanner in its home position, the focused beam expandsuntil it is collimated by the scan lens and then is re-focused by thecornea and lens of the patient onto the patient's retina. In thisposition, FDOCT A-scans of the patient's retina may be acquired formeasurement of the retinal position and axial reflectivity properties.Scanning of the galvanometer away from its central position results intranslation of the collimated beam away from the pupil center, where itwill eventually be clipped by the edge of the pupil. Thus, extensivelateral imaging of the retina may not be available in accordance withembodiments illustrated in FIGS. 28A and 28B.

As illustrated in FIG. 42, rapid insertion of the additional lens asdiscussed with respect to 28B for mode switching may be performed bymounting the lens or a sequence of identical lenses into a plate 4281,for example, alternating around the circumference of the plate with theabsence of a lens, which is then rapidly rotated into position by use ofa stepper or DC motor and suitable controller. In other words, thecontroller 4291 is configured to cause the mechanical means 4280, forexample, a stepper or DC motor, to rapidly rotate the plate 4281. Theplate 4281 includes both lenses 4280 and empty windows 4285, such thatas the plate 4281 is rotated, the system changes modes. It will beunderstood that embodiments are not limited to this configuration. Forexample, means for insertion of the lens could include mounting the lensin an arm attached to a rotary solenoid which could be rapidly rotatedinto and out of position without departing from the scope of the presentinventive concept.

Since the two dimensional optical scanner in these embodiments need onlydeviate a collimated or focused beam, the scanner clear aperture needonly be as large as the collimated beam size. In conventional systems,this collimated beam size may be less than about 5.0 mm, which enablesthe use of compact and high-speed galvanometer scanners.

It will be understood that switching between the anterior segment andretinal imaging modes as discussed with respect to FIGS. 28A and 28Balso involves changing the optical path length from the sample arm fibertip to the sample being imaged. Thus, changing of modes will alsorequire simultaneous re-setting of the reference arm position in commonFDOCT engine designs, which requires matching of optical path lengthbetween the sample and reference arms which will be discussed furtherbelow.

Referring now to FIGS. 29A and 29B, the system illustrated therein isconfigured for full two dimensional imaging of both the retina andstructures of the anterior segment. The retinal imaging system includesa collimating lens 2921, two two-dimensional galvanometer scanner pairs2910 and 2911, a scan lens 2902, and an objective lens 2901 placed asillustrated in FIGS. 29A and 29B. Those having skill in the art ofoptical design will realize that lenses with other focal lengths couldbe used to magnify or de-magnify the scan range and spot size on thepatient's retina and the working distance between the objective lens andthe patient's eye. The difference between this design discussed withrespect to FIGS. 28A and 28B is that the first three-dimensional scannerrequires both a large clear aperture and a large angular deviation toswitch between modes rather than insertion of an additional lens. Thesize requirement for the first two-dimensional scanner is that theaperture be as large as the desired scan range on the anterior segment(or suitably related to it if magnifying or de-magnifying optics areused in the telescope comprising the scan lens and objective lensfollowing the scanner), and that the scanner have sufficient angularexcursion to allow for switching to the anterior segment imaging mode byadjusting the correct angle required to hit the second two-dimensionalscanner pair. With the first scanner pair in this highly deviatedposition, the second two-dimensional scanner pair is used to image theretina with small angular excursions of the second galvanometer scannersperforming two-dimensional scanning of the focused sample arm beam onthe patient's retina. As illustrated in FIG. 29B, the secondtwo-dimensional optical scanner pair 2911 directs the re-directedcollimated light in a triangular pattern towards a curved mirror 2940placed on focal length f away from the large aperture original twodimensional scanner mirror pair. This alternative optical patheffectively transforms the scanning beam into a state such that theremaining optics along the optical path comprise a telescope whichimages the scanning focused beam into the anterior segment of thepatient's eye. Alternative embodiments of this alternative triangularoptical path may also be constructed along similar lines, for exampleusing a lens and flat mirror instead of a curved mirror, or usingmultiple curved mirrors in combination in order to reduce astigmatism.

In embodiments illustrated in FIGS. 29A and 29B, the optical path lengthof the optical system is longer in the anterior segment as compared tothe retinal imaging mode. If properly designed, this path lengthdifference could be designed to match the optical path length differenceanticipated upon traversing the length of a standard human eye,including the index of refraction of the acqueous and vitreous humor. Ifdone, this may eliminate the need for reference path length switchingwhen switching modes.

Further embodiments of systems configured to image both the retina andthe anterior segment of the eye 3060 will now be discussed with respectto FIGS. 30A and 30B. As illustrated, the imaging system includes acollimating lens 3021, two two-dimensional, a scan lens 3002, and anobjective lens 3001 placed as illustrated in FIGS. 29A and 29B. Asfurther illustrated in FIGS. 30A and 30B, embodiments illustrated inFIGS. 30A and 30B use a single two-dimensional scanner pair 3011 havinga large clear aperture and angular scan capability, and a flat mirror3041 in place of the second two-dimensional scanner pair discussed abovewith respect to embodiments illustrated in FIGS. 29A and 29B. Fortwo-dimensional scanning of the retina, the collimated beam entering thesample arm is incident on the two-dimensional scanner 3011, which isimaged into the pupil plane of the patient by a 4f or equivalenttelescope. Small angular deviations around the two-dimensional scannercenter position are imaged by the telescope into the patent's pupilplane. The patient's own cornea and lens act to focus this beam on theretina in a scanning pattern.

To switch to anterior segment imaging illustrated in FIG. 30B, thetwo-dimensional scanner pair 3011 is deviated by a large amount in orderto direct the incident collimated beam into a separate path consistingof a flat mirror 3041 and a concave mirror 3040 with focal length f, thelatter positioned a distance f from the two dimensional scanner 3011.Small deviations of the two-dimensional scanner about this large offsetdeviation now act to scan a focused beam across the surface of thetwo-dimensional scanner, which the 4f telescope then images onto theanterior segment of the patient's eye. If the 4f telescope is designedfor 1:1 imaging, then the clear aperture of the two-dimensional scannermust match the distance desired to be scanned on the patient's anteriorsegment. However, the telescope between the two-dimensional scanner pairand the eye may be alternatively designed to incorporate magnificationor demagnification as desired, albeit at the cost of additional scanangle requirements on the two-dimensional scanner.

Referring now to FIG. 31, a dual switchable reference delay for dualdepth imaging regions will be discussed. In the embodiments discussedabove with respect to FIGS. 28A through 30B for switching betweenanterior segment and retinal imaging, there is a need for equally rapidswitching of the FDOCT reference delay simultaneous with sample armoptics mode switching. Some embodiments for rapid switching areillustrated in FIG. 31. As illustrated therein, light from the referencearm of the interferometer 3151 is split into two or more separate pathsusing, for example, a fiber coupler 3152. The coupler is illustrated inFIG. 31 as a 2×2 coupler, however, it will be understood that higherorder couplers could also be used for rapid switching between more than2 reference delays (paths). In each delay arm 3171, 3172, a desiredoptical delay matching one of the modes of the sample arm scanner may bepre-set. To switch between reference delays, a rapid mechanical switch3153 may be used to block all but the desired reference delay. In someembodiments, the mechanical switch could be an arm mounted to a rotarysolenoid, a wheel with cutouts (akin to a chopper wheel) mounted to astepper or DC motor, or any other rapid mechanical switch familiar tothose familiar with the art of mechanical design without departing fromthe scope of the inventive concept.

Referring now to FIG. 32, a system for ocular spectral Domain OCTimaging will be discussed. As illustrated therein, the spectral domainOCT system includes a broadband optical source 3200, a comb filter 3201,a source path 3205, a beam splitter/combiner 3220, a reference path3207, a sample path 3208 with a scanning system and focal opticsillustrated in FIGS. 30A and 30B configured to appropriately to imagestructures of the sample, such as the cornea, anterior chamber, iris,lens, posterior chamber, and retina of the eye, a detector path 3206.The detector path 3206 includes a computer 3295 and a deep imaging,linear K Spectrometer 3232. As illustrated, the broadband source 3200 inFIG. 32 includes and SLD light source having a λ_(o) of about 850 nm anda Δλ of about 50 nm.

As further illustrated in FIG. 32, the reference path 3207 includes thefiber coupler 3270 as discussed above with respect to FIG. 31. The fibercoupler 3270 is configured to connect the reference path 3207 to atleast two other paths 3271, 3272 switched by a mechanical switch 3275,for example, solenoid or galvanometer. When the switch 3275 is inreference position 1, the system will operate in anterior segment modeand when the switch is in reference position 2, it will operate inretinal scanning mode.

Referring now to FIG. 33, a system for ocular swept source OCT imagingwill be discussed. As illustrated therein, the swept source OCT systemincludes a broadband optical source 3396, a source path 3305, a beamsplitter/combiner 3320, a reference path 3307, a sample path 3308 with ascanning system and focal optics illustrated in FIGS. 30A and 3013configured to appropriately to image structures of the sample, such asthe cornea, anterior chamber, iris, lens, posterior chamber, and retinaof the eye, and a detector path 3306. The detector path 3306 includes acomputer 3395, an A/D conversion switchable sampling rate 3399 and 500MHz photo-receiver 3398. As illustrated, the source 3396 in FIG. 33includes a 100 kHz Swept Laser source having a λ_(o) of about 1060 nmand a Δλ of about 100 nm.

As further illustrated in FIG. 33, the reference path 3307 a controller(AOM) 3360 as discussed above with respect to FIGS. 28A and 28B and 42.The controller 3360 is connected to a beamsplitter 3361, which isconfigured to split the light between first and second positions 3373,3374. The positions connect the reference path 3307 to at least twoother paths 3373, 3374 switched by dual position switches 3376 and 3377.When the switch is in reference position 1, the system will operate inretinal imaging mode 3373 and when the switch is in reference position2, the system will operate in anterior imaging mode 3374.

As discussed above, for comprehensive FDOCT imaging of the eye byrapidly switching between imaging modes designed for imaging differenceocular structures along the visual axis, it would be desirable for theimaging depth (axial field of view) of each mode to be optimized for theexpected length and desired axial sampling density of each structure.For example, for imaging of the entire anterior segment, the optimalimaging depth is the expected maximum anterior segment depth of theanticipated patient population, which may be as long as 6-8 millimeters.For imaging of the retina, which is less than 1.0 mm thick in mostlocations and contains many closely spaced layers and structures, it maybe preferable for the retinal imaging mode to have a shorter imagingdepth and denser sampling within it.

As discussed above, in all FDOCT systems, there is an inverserelationship between the imaging depth z_(max) and the spectral samplinginterval in wavenumber units δ_(s)k given by:

$\begin{matrix}{z_{\max} = {\frac{\pi}{{2 \cdot \delta_{s}}k}.}} & \left( {3\mspace{14mu} {above}} \right)\end{matrix}$

The total sampled spectral width is given by the spectral samplinginterval δ_(s)k multiplied by the number of spectral samples acquiredper A-scan, typically several thousand, and thus the depth samplingdensity is given by the imaging depth divided by the number of spectralsamples (or some multiple of that number if interpolation is performed).In SDOCT systems, the spectral sampling interval δ_(s)k is typicallyfixed by the spacing of the pixels on the array detector used in thespectrometer and the magnification and spectral dispersion of theinternal optical elements of the spectrometer. In SSOCT systems,however, the spectral sampling interval δ_(s)k is determined by thesweep rate of the light source and/or the electronic sampling rate ofthe analog to digital converter which is recording the SSOCT signal, atleast one of which may be rapidly adjustable electronically or by othermeans. In the case of SSOCT, therefore, it will be desirable to adjustthe spectral sampling interval and thus the imaging depth and depthsampling density (according to the prescription in Equation 3) on thefly according to the structure or part of the eye which is being imaged.This imaging depth switching may be coupled to the sample and referencearm mode switching embodiments described above with respect to FIGS. 28Athrough 33, such that when switching the sample arm optics and referencearm delay from the anterior segment to the retina, for example, theimaging depth is also switched to allow for optimal imaging depth andsampling density of retinal structures. In some embodiments, the imagingdepth and depth sampling density may be varied within a single mode ofthe sample and reference arm optics, for example to switch between shortimaging depth, high spatial sampling density imaging of the cornea andlong imaging depth, lower spatial sampling density imaging of the entireanterior segment.

FDOCT systems exhibit a decrease in signal-to-noise ratio (SNR) as afunction of path length difference between the sample and reference arms(and thus the distance from the origin in FDOCT images), which isrelated to the spectral resolution of the FDOCT system, δ_(r)k Rapidsensitivity falloff is a drawback in FDOCT systems because it limits theamount of the imaging depth which actually contains useful imageinformation. The sensitivity “falloff” may be characterized by theimaging depth at which the sensitivity falls to 6 decibels below itsvalue at the zero path length difference location. This value isinversely related to the system spectral resolution δ_(r)k:

$\begin{matrix}{{\hat{z}}_{6d\; B} = {\frac{2{\ln (2)}}{\delta_{r}k}.}} & \left( {4\mspace{14mu} {above}} \right)\end{matrix}$

In unmodified SDOCT systems, δ_(r)k is usually limited by the spectralresolution of the spectrometer (including the finite spacing of the CCDpixels and diffraction in the spectrometer). In unaltered SSOCT systems,δ_(r)k is typically limited by the instantaneous lineshape of the sweptlaser source, although other factors such as the bandwidth of thedetection electronics may also come into play.

As discussed above, conventional methods exist for decreasing SNRfalloff in FDOCT systems by introducing a comb filter into the FDOCTsystem (either in the source arm, both sample and reference arms, ordetector arm), such that the spectral extent of light collected at eachspectral sampling interval δ_(s)k is limited by the transmissioncharacteristics of the comb filter rather than the spectral resolutionof the spectrometer (in SDOCT) or the instantaneous linewidth of theswept laser source (in SSOCT) (U.S. Pat. No. 7,602,500). Such a combfilter may be implemented as a Fabry-Perot etalon or filter, having afree spectral range (FSR) set to be equal or nearly equal to the desiredFDOCT spectral sampling interval δ_(s)k, and a full width athalf-maximum (FWHM) transmission peak width set to be equal or nearlyequal to the desired FDOCT spectral resolution δ_(r)k required toachieve a given 6 dB falloff length {circumflex over (z)}_(6dB)according to the formula in Eq. (4). Thus, the comb filter willessentially modify the spectrum reaching the FDOCT detector such thatthe optical bandwidth detected at each spectral sampling interval isdecreased, thus decreasing SNR falloff.

In comprehensive ocular SSOCT systems as discussed above wherein thespectral sampling interval and depth sampling density are adjusted asper Equation 3 according to the structure or part of the eye which isbeing imaged, it is desirable to further implement a comb filter fordecreasing the extent of SNR falloff which is also suitably adjustableto maintain the comb spacing or FSR as the spectral sampling interval isadjusted. In Fabry-Perot etalons, the FSR is related to the thickness ofthe etalon, the index of refraction of the material inside the etalon,and the angle of light incidence upon the etalon. According to someembodiments, one or more of these parameters should be varied insynchrony with changing the spectral sampling interval δ_(s)k in orderto keep the comb filter peaks within their respective spectral samplingintervals. In some embodiments, this may be done by employing aso-called Fabry-Perot tunable filter, which utilizes a piezo-electricelement to electronically tune its FSR. Electronic control of the FSR ofsuch a filter may be electronically coupled to the mechanism forchanging the spectral sampling interval δ_(s)k, for example by changingthe digitization rate of the analog-to-digital converter.

Several methods exist in the prior art for increasing the imaging depthz_(max) by a factor of two by resolving the so-called “complexconjugate” or “mirror image” artifact in FDOCT, which not only limitsthe maximum imaging depth for a give spectral sampling interval but alsointroduces unwanted additional image artifacts. These prior art methodsinclude techniques borrowed from phase shift interferometry involvingmultiple sequential or simultaneous A-scan acquisitions with referencepath delays varying by a multiple of pi/2 radians.

For an SDOCT system designed for comprehensive ocular imaging accordingto all of the embodiments of the present inventive concept, thepreferred embodiment for complex conjugate removal (CCR) is viasinusoidal phase modulation as discussed in U.S. Pat. No. 7,742,174,Sinusoidal phase modulation involves placement of a sinusoidal pathlength modulation in either the sample or reference arm of an SDOCTsystem which varies the differential path length between the arms withamplitude and phase given in the text preceding Equation (14) in U.S.Pat. No. 7,742,174, at a rate corresponding to Π/4 radians of sinusoidalmodulation per A-scan integration time of the spectrometer. Then, eachset of four sequential A-scan acquisitions are combined according toEquation (14) of U.S. Pat. No. 7,742,174 in order to generate an A-scanwith total depth equal to 2*z_(max) as defined above. If the amplitude,phase and frequency of the sinusoidal modulation are set exactly asspecified in U.S. Pat. No. 7,742,174, then the resulting A-scan shouldtheoretically be completely free of DC, auto correlation, and complexconjugate artifacts. However, slight deviations from perfection inachieving these parameters such as will be experienced in any realphysical implementation of sinusoidal phase modulation may lead to adegradation of performance compared to the ideal result in the form ofincomplete complex conjugate artifact suppression. Thus, an additionalstep of applying quadrature projection processing according to FIG. 2 ofU.S. Patent Application Serial No. 2008/0170219 may be applied toimprove the complex conjugate artifact rejection, at the cost of a smallamount of reduced signal to noise ratio. Quadrature projectionprocessing is an algorithmic step which does not require any hardwaremodification and which reduces the complex conjugate artifact fromimperfectly phase modulated SDOCT data by forcing the real and imaginaryparts of the recorded A-scan signal to be orthogonal.

For an SSOCT system designed for comprehensive ocular imaging accordingto all of the embodiments of the present inventive concept, thepreferred embodiment for complex conjugate removal (CCR) is theso-called “heterodyne” CCR method, which involves introducing afrequency shift between the sample and reference arm light and thusshifting the carrier frequency of the image-bearing signal away from DC,about which the complex conjugate artifact is centered as discussed inU.S. Pat. No. 7,336,366. With the addition of this frequency shift, theA-scan free of complex conjugate artifact is found from the Fouriertransform of the detected signal, centered at the frequency shift value.If an A/D converter is used which has much higher bandwidth than theSSOCT signal itself, then the frequency shift value can be set to bemany times the frequency encoding the z_(max) value of the A-scan, thusthe complex conjugate artifact will be located far in frequency spaceaway from the A-scan data. If a very high sweep speed is employed,however, such that the SSOCT signal already occupies a substantialfraction of the A/D converter bandwidth, then the complex conjugateartifact may only be shifted to the borders of the depth-doubled A-scan.This method of heterodyne CCR is consistent and will not interfere withthe embodiments described above for switching between sample andreference arm imaging modes, switching SSOCT imaging depth, andswitching of the comb filter FSR spacing to remain consistent with thespectral sampling interval.

Referring now to FIGS. 34 through 38, further embodiments of systemsconfigured to switch between scanning modes will be discussed. Asdiscussed above, for whole Eye imaging, two distinct methods of scanningtypically must be used. The cornea, anterior chamber and lens scanningrequires a telecentric type scanning which, as discussed above, can bedefined as each filed point or angle of the scanning mirror producesfocus rays which are parallel to the optical axis. Retinal scanning, onthe other hand, typically requires that a collimated beam is impingingon the cornea with the conjugate of the scanning mirror be located atthe pupil of the eye. For biometry, collimated light must be impingingon the cornea however since only axial scans are required the conjugateof the scanning mirror does not need to be placed at the pupil of theeye. Conventional OCT systems require separate optics to do each type ofscan. The challenge to doing either telecentric scanning plus axialretinal scanning for biometry or telecentric scanning plus retinalscanning for true dual scan modes is developing a method to rapidlychange the imaging optics or rapidly alter the characteristics of theimaging optics.

Various embodiments for switching between modes will now be discussedwith respect to FIGS. 34 through 38. In addition to switching techniquesto change from concentric imaging of the retina to telecentric imagingof the cornea, a zoom lens type optical translation can be used to movebetween the two modes of operation. The advantage to these embodimentsis that it affords the ability to focus the scanning beam at differentpoints in the eye for the highest lateral resolution.

Referring first to FIG. 34, an optical layout for telecentric scanningmode will be discussed. As illustrated, the system of FIG. 34 forimaging an eye 3460 includes an XY scanning mirror 3457, a telecentricscanning lens 3423, a fiber conjugate plane 3425, first and secondobjective lenses 3435 and 3436, XY scanning mirror conjugate 3458 and afiber conjugate 3459. The telecentric scanning lens 3423 forms aconjugate of the input fiber (not shown) which is an intermediatetelecentric image plane. Therefore, the focal length of the telecentricscanning lens is totally independent of the objective focal length andis selected for both optimum mechanical layout and as well as overallsystem numerical aperture. The combined focal length of the objective isselected to give the widest possible field of view at the desiredworking distance or distance from the last optical element to the corneaof the eye. As used herein, f₁ is equal to the focal length of thetelecentric scanning lens and f₂ is equal to the sum of the focallengths of objective lens 1 and objective lens 2.

Referring now to FIG. 35, the system of FIG. 35 for imaging an eye 3560includes an XY scanning mirror 3557, a telecentric scanning lens 3523, afiber conjugate plane 3525, combined first and second objective lenses3537, XY scanning mirror conjugate 3558 and a fiber conjugate 3559. Theoptical system of FIG. 35 is in the Concentric or Retinal scanning mode.As illustrated, in this configuration the objectives lenses 3537 arecombined by bringing them into close proximity to one another. Thiscombines the optical power of the objective lenses and moves thescanning mirror conjugate to the cornea and thus the fiber is imagedonto the retina. As compared to the spacing in the telecentric mode, thetravel of objective lenses is the differential between the two settingsby the factors shown in FIG. 34. The working distance or the distancefrom the last optical element to the eye however remains fixed.

In FIG. 35, f₁ is equal to the focal length of the telecentric scanninglens 3523 and f₂ is equal to the sum of the focal lengths of objectivelens 1 and 2 3537. The OCT reference arm position typically tracks thelens translation. Other combinations of focal lengths and spacing can beused without departing from the scope of embodiments discussed herein.Thus, as illustrated in FIGS. 34 and 35, objective lens A 3435 slides tochange from iris pivot mode (object lens A 3435 proximate objective lensB 3436) to telecentric mode (objective lens A 3435 forms relay toobjective lens B 3436).

Referring now to FIG. 36, a system in telecentric mode for imaging thecornea of the eye 3660 includes a fiber input 3676, a collimating lens3678, a scanning mirror 3677, a telecentric scanning lens 3623, and atelecentric scanning beam. For the case of biometry, a concentric scanis not necessary and changing the scan mode from telecentric tocollimated will allow the OCT beam to be focused onto the retina. Tochange the system of FIG. 36 from telecentric to collimated, thecollimating lens 3778 can be translated by a distance equal to its focallength as illustrated in FIG. 37. The system in FIG. 37 includes acollimating lens 3778, a scanning mirror 3777, a telecentric scanninglens 3723, a collimated beam 3787 and a fiber conjugate 3788.

In some embodiments, to change the system of FIG. 36 from telecentric tocollimates, a secondary lens 3879 can be inserted behind the collimatinglens 3878 to change the focal position and illustrated in FIG. 38. Thesystem of FIG. 38 includes a collimating lens 3878, a scanning mirror3877, a telecentric scanning lens 3823, a collimated beam 3887 and afiber conjugate 3888.

Some specific embodiments will now be discussed. To achieveaccommodation for both myopic and hyperopic the lens set must be shiftedfrom the conjugate plane with an offset of 2.0 mm for a +12 diopteraccommodation and −1.75 mm for a −12 diopter accommodation. The lenspositions and spacing for telecentric imaging mode however remainconstant. Therefore the translation mechanism allows for the translationof the lens pair when moved into the retinal position.

To translate the lenses from telecentric to retinal mode imaging, astandard zoom lens double helix drive can be incorporated for bothmanual and automated actuation. Automated means can be accomplishedwith, for example, stepper motors, piezo motor, solenoids or voicecoils, but are not limited thereto. With proper mechanical coupling,each methodology has the capacity to switch modes well within a singlesecond. Both stepper and piezo motor drives afford the ability to addprogrammability to the lens translation allow intermediate surfaces tobe imaged at a high lateral resolution.

Due to the physical property of diffraction, high resolution scanninglenses are constrained to have low depth of focus resulting indecreasing lateral resolution as the distance from the image plane isincreased. To alleviate this issue long depth of focus optics can bedesigned but at the sacrifice of resolution. The depth of focus (d) isdefined as:

d=2πω_(o) ²/λ  (38)

where ω_(o) is the Airy radius which is the radius of the firstdiffraction minimum of the focuses spot and λ is the wavelength oflight. Therefore as the relation shows, with the wavelength fixed, thelarger the Airy radius the greater the depth of field and since the Airyradius also defines the scanning resolution, the lower the resolution.

Likewise, it is possible to derive the required Airy radius based uponthe desired scan depth as follows:

ω_(o) √dλ/2π  (39)

In embodiments where a desired scan depth for the cornea to lens is adistance of 6.55 mm, the scanning resolution will be limited to 29.6 μm.This can be achieved by simply changing the focal length of thecollimator used in system proportionally to the focal length of thescanning optics and making no other changes to the scanning optics fromthe current product offering. No additional optics design work isrequired. Since the current telecentric scanning optics have an Airydisk radius of 12 μm a 2.5× reduction in the focal length of thecollimator will produce the 29.6 μm Airy radius required.

For volume phase holographic grating based spectrometer design, theimaging depth is related to the dispersion characteristics of thespectrometer in the following expression:

Z=λ _(c) ²/4nλ _(s)  (40)

Where

-   -   δ=source bandwidth (nm)    -   λ_(c)=source center wavelength (nm)    -   p=pixels (detector channels)    -   λ_(s)=spectrometer wavelength spacing=δ/p    -   n=index of refraction        Therefore in order to design the spectrometer for the maximum        desired imaging window depth of 6.55 mm for the region of cornea        to the posterior surface of the lens the above equation is        solved for

With λ_(s)=δ/p equation (1) becomes

Z=pλ _(c) ²/4nδ  (41)

Given that p is determined by available detectors and therefore is afixed value solving for in becomes,

δ=pλ _(c) ²/4nZ  (42)

Alternatively, setting δ to a known value and solving for X innanometers leads to,

λ_(c)=√(4nδZ/p)  (43)

The determination of the optimum values for λ_(c) and δ are based uponthe design models for the source.

Further definition of the design parameters can be obtained by relatingthe image size to the detector pixel size in order to determine thespectrometer focal length required. Assuming a collimated beam input tothe grating the diffraction limited spot size can be represented by thefollowing expression:

D=1.22λ_(c)(f/d)  (44)

Where f is the focal length of the spectrometer imaging optics and d istypically the lens aperture diameter which in this case is equivalent tothe spectrometer input collimated beam diameter. Solving for (f/d),

(f/d)=D/1.22λ_(c)  (45)

Therefore given a pixel size of 10 μm and setting the target diffractionlimited image spot radius to the detector pixel size with a 75% fillfactor as is standard practice yields a spot diameter of 7.5 μm. FromEquation 45, we can determine the ratio of the focal to input beamdiameter,

(f/d)=3.5  (46)

From expression (46) for a collimated beam of 25 mm in diameter therequired focal length of the spectrometer imaging optics is 89 mm.Conversely, setting the focal length to 100 mm requires a 28 mmcollimated beam input. The determination of which parameter to solve foris based solely on the design constraints of the spectrometer.

Using the grating equation:

λ_(c) f=sin θ_(i)+sin θ_(d)  (47)

Where

-   -   λ_(c)=source center wavelength    -   f=spatial frequency of the grating    -   θ_(i)=angle of incidence    -   θ_(d)=angle of diffraction

For VPH grating designs θ_(i)=θ_(d), Solving for f equation 5 is reducedto:

f=2 sin θ/λ_(c)  (48)

With the practical upper limit established by:

f=2/λ_(c)  (49)

Since the dispersion efficiency of the VPH grating is inverselyproportional to the spatial frequency design optimization is directedtoward reducing the spatial frequency. The optical design of thespectrometer is also critical in selecting the grating dispersion value.Since the array has a predetermined physical length and the centerwavelength and bandwidth are fixed by the desired imaging depth, thedispersion is selected to insure full coverage of the spectral bandwidthacross the detector array.

By definition, the dispersion of the grating is the rate of change ofthe angle of diffraction with wavelength for a fixed angle of incidenceor Δθ/Δλ which from a differentiation of equation 47 yields:

Δθ/Δλ=f/cos θ  (50)

The dispersion of the grating is also related to the required geometryof the spectrometer optics. For a given array length and focal length ofthe imaging optics the angle of dispersion can be given as:

Φ=2 tan⁻¹(A/2f)  (51)

where f is the focal length of the imaging optics and A is the detectorarray length.

From equation (50) the grating dispersion relates the dispersion angleby:

Φ=afδ/cos θ=2 tan⁻¹(A/2f)  (52)

where a is the unit conversion from radians/mm to degrees/nm and f isdetermined by the detector pixel size as stated in equation 46.

The expression can be reduced as follows:

(aδ/λ _(c))tan θ=tan⁻¹(A/2f)  (53)

Solving for θ:

θ=tan⁻¹[(λ_(c) /aδ)tan⁻¹(A/2f)]  (54)

From the above equations, the required dispersion angle can becalculated for a given spectrometer layout. The parameters required asinputs to the equations are the detector pixel size which defines therequired focal length, (A) the linear dimension of the detector array,(λ_(c)) the center wavelength of the source and (δ), the bandwidth ofthe source. From the calculated dispersion value, the grating frequencyand grating angle can be calculated resulting in a completecharacterization of the spectrometer design.

Various methods for whole eye imaging are illustrated in the flowchartsof FIGS. 39-41. In particular, referring first to FIG. 39, methods forhigh resolution whole eye imaging using a 3.5 mm single-sided windowwith depth doubling will be discussed. As illustrated in FIG. 39, aregion under test may be selected may be selected in block 3900. Asillustrated, the region under test 3900 may be the cornea 3901, theanterior segment 3902, the crystalline lens segment 3903, the posteriorsegment 3904 or the retina 3905. It will be understood that embodimentsof the present inventive concept are not limited to the regions in box3900 and that more or less regions may be enumerated without departingfrom the scope of the present inventive concept.

Once the region is selected (block 3900) the FDOCT settings 3910 aredetermined. For example, in some embodiments z_(max) is set to 3.5 mm(block 3911). At this point if the cornea 3901 is the selected region,the reference is set to anterior to corneal apex (block 3912) and a 3.5mm image may be acquired (block 3913). If the anterior segment isselected (block 3902), the reference is set to middle of anteriorsegment (block 3914) and a 7.0 mm image may be acquired (block 3915). Ifthe crystalline lens segment is selected (block 3903), the reference isset to middle of lens (block 3916) and a 7.0 mm image may be acquired(block 3917). If the posterior segment is selected (block 3904), thereference is set to middle of posterior chamber (block 3918) and a 7.0mm image may be acquired (block 3919). If the retina is selected (block3905), the reference is set to anterior to retina (block 3920) and a 3.5mm image may be acquired (block 3921).

As further illustrated in FIG. 39, the CCR setting (3930) is set to“off” if the cornea (3901) or the retina (3905) is selected as theregion under test (3900) and set to “on” if the anterior segment (3902)is selected as the region under test (3900).

Finally, as further illustrated in FIG. 39, scan options (3940) may alsobe selected/set. For example, if the cornea 3901 is the selected region,telecentric optics are selected focused in the vicinity of the cornea(block 3941). If the anterior segment is selected (block 3902), thefocus of the optics is shifted towards center of the anterior segment(block 3942). If the crystalline lens segment is selected (block 3903),the focus of the optics may be shifted towards the center of the lens(block 3943). If the posterior segment is selected (block 3904),iris-pivot optics are selected with focus on mid-posterior chamber(block 3944). If the retina is selected (block 3905), the focus isshifted to the retina (block 3945).

Referring now to FIG. 40, methods for three-step whole eye imaging forBiometry using 7.0 mm single-sided window with depth doubling andquasi-telecentric optics will be discussed. As illustrated in FIG. 40, aregion under test may be selected may be selected in block 4000. Asillustrated, the region under test 4000 may be anterior chamber pluslens 4001, the posterior lens capsule plus posterior chamber 4001 or theposterior view to capture retina 4003. It will be understood thatembodiments of the present inventive concept are not limited to theregions in box 4000 and that more or less regions may be enumeratedwithout departing from the scope of the present inventive concept.

Once the region is selected (block 4000) the FDOCT settings 4010 aredetermined. For example, in some embodiments z_(max) is set to 7.0 mm(block 4011). At this point if the anterior chamber plus lens 4001 isthe selected region, the reference is set to bottom of anterior chamber(block 4012) and a 14 mm image may be acquired (block 4013). If theposterior lens capsule plus posterior chamber is selected (block 4002),the reference is set to mid-way into target of posterior (block 4014)and a 14.0 mm image may be acquired (block 4015). If the posterior viewto capture retina is selected (block 4003), the reference is set totowards inner retinal surface (block 4016) and a 14 mm image may beacquired (block 4017).

As further illustrated in FIG. 40, the CCR setting (4030) is set to “on”if the anterior chamber plus lens (4001) is selected as the region undertest (4000).

Finally, as further illustrated in FIG. 40, scan options (4040) may alsobe selected/set. For example, if the anterior chamber plus lens (4001)is the selected region, telecentric optics are selected focused towardsa bottom of the anterior chamber (block 4041). If the posterior lenscapsule plus posterior chamber is selected (block 4002), the focus ofthe optics is shifted midway into target of the posterior (block 4042).If the posterior view to capture retina is selected (block 4003), thefocus of the optics may be shifted towards the retinal surface (block4043).

Referring now to FIG. 41, methods for variable range whole eye imagingusing a 7.0 mm single-sided window with depth doubling will bediscussed. As illustrated in FIG. 41, a region under test may beselected may be selected in block 4100. As illustrated, the region undertest 4100 may be the anterior chamber 4101, the anterior chamber pluslens 4102, the crystalline lens 4003, the posterior lens plus posteriorchamber 4104, posterior segment 4105, posterior view to capture retina4106 or retina and choroid 4107. It will be understood that embodimentsof the present inventive concept are not limited to the regions in box4100 and that more or less regions may be enumerated without departingfrom the scope of the present inventive concept.

Once the region is selected (block 4100) the FDOCT settings 4110 aredetermined. For example, in some embodiments z_(max) is set to 7.0 mm(block 4111). At this point if the anterior chamber 4001 is the selectedregion, the reference is set to interior to anterior chamber (block4112) and a 7.0 mm image may be acquired (block 4113). If the anteriorchamber plus lens (block 4102) is selected, the reference is set tobottom of anterior chamber (block 4114) and a 14.0 mm image may beacquired (block 4115). If the crystalline lens is selected (block 4103),the reference is set anterior to lens (block 4116) and a 7.0 mm imagemay be acquired (block 4117). If the posterior lens plus posteriorchamber is selected (block 4104), the reference is set to mid-lens(block 4118) and a 7.0 mm image may be acquired (block 4119). If theextended range and field of view for posterior visualization is selected(block 4105), the reference is set to midway into target of posterior(block 4120) and a 14 mm image may be acquired (block 4121). If theposterior view to capture retina is selected (block 4106), the referenceis set to towards inner retinal surface (block 4122) and a 14 mm imagemay be acquired (block 4123). If the optimized range for outer vitreous,retina and choroid is selected (block 4107), the reference is set totowards inner retinal surface (block 4124) and a 7.0 mm image may beacquired (block 4125).

As further illustrated in FIG. 41, the CCR setting (4130) is set to“off” if the anterior chamber (4101), crystalline lens (4103) or theoptimized range for outer vitreous, retina and choroid is selected(block 4107) is selected as the region under test (4100) and set to “on”if the anterior chamber plus lens (4102) or extended range and field ofview for posterior visualization is selected (block 4105) is selected asthe region under test (4100).

Finally, as further illustrated in FIG. 41, scan options (4140) may alsobe selected/set. For example, if the anterior chamber (4101) is theselected region, the optics may be focused interior to the anteriorchamber (block 4141). If the anterior chamber plus lens is selected(block 4102), the focus of the optics is shifted towards a bottom of theanterior chamber (block 4142). If the crystalline lens is selected(block 4103), the focus of the optics may be shifted towards the centerof the crystalline lens (block 4143). If the posterior lens plusposterior chamber is selected (block 4104), optics are focused onposterior lens capsule (block 4144). If the extended range and field ofview for posterior visualization is selected (block 4105), theiris-pivot focus is set to mid-posterior chamber (block 4145). If theposterior view to capture retina is selected (block 4106), the focus isset towards the retinal surface (block 4145). If the optimized range forouter vitreous, retina and choroid is selected (block 4107), the focusis shifted to optimize the retina (block 4147).

In the drawings and specification, there have been disclosed exemplaryembodiments of the present inventive concept. However, many variationsand modifications can be made to these embodiments without substantiallydeparting from the principles of the present inventive concept.Accordingly, although specific terms are used, they are used in ageneric and descriptive sense only and not for purposes of limitation,the scope of the inventive concept being defined by the followingclaims.

1. (canceled)
 2. An optical coherence tomography system for imaging awhole eye, the system comprising: a sample arm including focal opticsconfigured to switch between at least two scanning modes, wherein the atleast two scanning modes comprise an anterior segment scanning mode anda retinal scanning mode; wherein the sample arm of the system in retinalscanning mode comprises a collimator, a two-axis scanning mirrorassembly, and a scan lens assembly, an objective lens; and wherein thefocal optics of the sample arm are configured to reposition at least onemovable lens to change an optical coherence tomography (OCT) scan beamfrom a collimated beam for imaging an emmetropic eye to one of adiverging beam for imaging a myopic eye and a converging beam forimaging a hyperopic eye; a mechanical means for modifying componentswithin an optical pathway of the sample arm to switch the scanning modebetween the retinal scanning mode and the anterior segment scanningmode; and a reference arm coupled to the sample arm and comprising ameans for discrete switching between two path delays in coordinationwith switching between the at least two scanning modes, wherein a firstpath delay is configured to match an optical path length of the OCT scanbeam in the retinal scanning mode and a second path delay is configuredto match an optical path length of the OCT scan beam in anterior segmentscanning mode.
 3. The system of claim 2, wherein the focal optics of thesample arm are further configured to reposition the at least one movablelens using a mechanical means for insertion of an additional lensassembly into the optical pathway of the sample arm immediately proximalor immediately distal to the collimator to change the system from theretinal scanning mode to anterior segment scanning mode.
 4. The systemof claim 2: wherein the sample arm of the system in the anterior segmentscanning mode comprises a collimating lens, two two-axis scanning mirrorassemblies, a scan lens, an objective lens and a curved mirror placed afocal length (f) away from a first of the two two-axis scanning mirrorassemblies, wherein a second of the two two-axis scanning mirrorassemblies is configured to direct re-directed collimated light in atriangular pattern towards the curved mirror causing the optical pathlength of the system to be longer in the anterior segment scanning modethan in the retinal scanning mode.
 5. The system of claim 2: wherein thesample arm of the system in the anterior segment scanning mode comprisesa collimating lens, a two-axis scanning mirror assembly, a scan lens, anobjective lens, a flat mirror and a concave mirror placed a focal length(f) away from the two-axis scanning mirror assembly, wherein lightincident on the two-dimensional scanner pair is deviated such that thean incident collimated beam is directed into a separate path consistingof the flat mirror and the concave mirror.
 6. The system of claim 2,wherein the means for discrete switching a reference delay comprises arapid mechanical switch configured to block all but a identifiedreference delay associated with a corresponding one of the at least twoscanning modes.